X-ray tomography

ABSTRACT

An x-ray tomography system which can generate a qualitative 3D image of a region of interest using a an x-ray source, the x-ray source configured to emit x-ray radiation at the region of interest. The x-ray radiation or the x-ray source or the relative position of the x ray source configured to be moved in a two dimensional plane. An x-ray detector including a plurality of detector elements arranged in a two dimensional plane opposite the x-ray source, the x-ray detector configured to detect x-ray radiation after attenuation by the subject and provide an indication of the detected x-rays. And a processor configured to receive the indication of the detected x-rays and resolve the detected x-ray radiation into a three dimensional image. The three dimensional image is qualitative in nature.

INCORPORATION BY REFERENCE TO ANY PRIORITY APPLICATIONS

Any and all applications for which a foreign or domestic priority claimis identified in the Application Data Sheet as filed with the presentapplication are hereby incorporated by reference under 37 CFR 1.57.

This application is a continuation of U.S. patent application Ser. No.17/024,467, filed Sep. 17, 2020, titled X-RAY TOMOGRAPHY,” which claimsthe benefit under 35 U.S.C. § 120 and 35 U.S.C. § 365(c) as acontinuation of International Application No. PCT/US2019/022820,designating the United States, with an international filing date of Mar.18, 2019, titled “X-RAY TOMOGRAPHY,” which claims the benefit of U.S.Provisional Patent Application No. 62/645,163, filed Mar. 19, 2018; U.S.Provisional Patent Application No. 62/677,312, filed May 29, 2018; U.S.Provisional Patent Application No. 62/711,522, filed Jul. 28, 2018; andU.S. Provisional Patent Application No. 62/729,433, filed Sep. 11, 2018.The entirety of each of the aforementioned applications is incorporatedby reference herein.

TECHNICAL FIELD

The present disclosure relates three-dimensional x-ray imaging formedical and industrial applications.

BACKGROUND

Three-dimensional (3D) x-ray images are typically generated usingcomputed tomography (CT), comprised of an x-ray source generating a fanbeam, or point beam, and a linear or corresponding point detector orsometimes a small format 2D detector which is arc shaped. Cone Beam CT,a variation of CT, typically comprises a C-arm-mounted CT unit and adigital flat panel detector where either the subject or thesource/detector rotates about an axis. Multiple images need to be takenof the entire subject in large angles, typically more than 180° toreconstruct a 3D image of the subject. This means that the subjectreceives a relatively high dosage of radiation. This process is alsotime-consuming. Present 3D CT systems are not suitable for portability,especially outside of hospitals or surgical centers or mobilediagnostics and surgical stations.

Presently non-rotating tomography is typically achieved by using one ormore 2D flat panel detector(s) combined with two or more x ray sourceson a linear array or an arc array or a 2D plane or 3D space.

Inverse geometry CT tomography-based techniques requires the use of atwo-dimensional (2D) collimator with holes combined with a scanningx-ray source that emits through these holes before passing through thesubject.

The resultant tomography images in non-rotation CT systems are not highresolution and cannot offer quantitative information generally providedby a 3D CT scanner.

Scatter interference is a major issue in both rotational andnon-rotational CT.

In rotational CT, Scatter to primary x ray Ratio (SPR) scales almostproportional to the rotating axis angle coverage of the X ray cone beam.However, with lower rotation axis coverage, longer scanning times arerequired for images acquisition needed for tomography.

Currently, major efforts are spent on new concepts for Spectral CT,which uses the spectral information of transmitted X-rays to extractadditional information about the scanned patient or object. SPR is highin Spectral CT. For lower energies, the primary intensity isincreasingly covered by scattered radiation. For the constraint that thescatter-to-primary ratio is recommended to not exceed a value of one,the measured energies of single photons are recommended to be employedfor spectral processing above a certain limit, for example,approximately between 30-35 KeV.

In CT, with increasing scatter to primary ratio (SPR, image artifactsemerge and degradation in image quality.

Anti Scatter Grid (ASG) is a key solution for reducing scatter in CT.Even with ASG corrections, the SPR remain to be for example, 20% or muchmore for many energy levels.

For high resolution imaging, Contrast to Noise Ratio, CNR an indicatorfor image quality, may not improve significantly and the visibility oflow contrast details may be reduced with ASG.

SUMMARY

Most current CT reconstruction theory assumes that X-ray photons areabsorbed or pass through the subject illuminated without interaction,meaning when SPR is 0. In human body imaging, the ratio of scatter tothe primary signal is generally as high as between 50% and 100%. Thepresent disclosure provides a system which can generate a complete 3Dimage of a region of interest from a two dimensional scanning process.The generated image can have a thickness or depth, or dimension (forexample, in a z axis plane) comprising three or more datapoints. Eachdata point can be resolvable by x-ray measurement or opticalmeasurements or other dimensional measurement methods. This can be doneby minimizing the number of measurements using a 2D detector. This canbe done by minimizing introduction of unknowns pixels outside of theregion of interest in 1D, 2D and 3D space. Minimizing the introductionof unknowns can be done by minimizing the movement distances andmovement dimensions, by minimizing hardware and movement complexity, byminimizing x-ray acquisition time, by minimizing radiation exposureand/or by simplification of 3D x-ray imaging method and apparatus ofprior art.

By using a compact system with a small number or no moving parts, thisdisclosure provides a three dimensional x-ray imaging system using a twodimensional system. The system can also generate x ray measurements fora region of interest of a subject with multiple components internally,which can be analyzed quantitatively including material decomposition,identification and characterization and determination measurements ofphysical and chemical characteristics of the components, such as atomicweight, density, detailed dimensions and microstructures and features inspace, localization of components in space, dynamic movementscharacteristics, fluid dynamics, temporal marker identification andlocalization, flexibility of the components, interaction betweencomponents, such as molecular and/or cellular interactions,identification of a component based on one of more such characteristics.The system can provide all of the above benefits in one system. Thistype of system has wide application in hospitals, surgical centers, andmobile stations. It also provides portability of such a system inconventional portable formats, carry-on, foldable for field applicationssuch as sports medicine, veterinary, remote and ambulatory diagnosticsand imaging guidance and material identification in the field. Thepresent disclosure can also provide ultrafast image construction of 3Dimages to enable real time measurement and time dependent measurementsin multiple dimensional imaging.

The present disclosure can also enable quantitative imaging for othernon-rotational tomography systems to have more capabilities, for exampleInverse Geometry Scanning fluoroscope, multiple dimensional system using2D detectors coupled with motorized x-ray sources which can move in anarc or linearly, and/or static multiple x-ray sources, systems with Xray source(s) can place or move in 1D, 2D or 3D dimensions, pixelated 2Dflat panel x-ray sources, or an x ray source comprised of a cathode,each having a plurality of individually programmable electron emittingunits for emitting an electron beam when an electric field is applied,an anode target that emits X-ray beam when subjected to collision of theelectron beams emitted an X-ray source, and/or metal liquid jet sources.

The present disclosure improves imaging modalities such as x raymicroscopes and interferometry including Fourier transform methods,x-ray interferometry, coherence x-ray phase contrast x ray imaging,coherence contrast x-ray imaging, by providing a fast and non-rotationalCT system and in some cases, providing quantitative capabilities.

The speed bottle neck in acquiring fast 3D images in the prior art isgenerally the speed of the motion system, compared to the presentdisclosure, the speed limiting factor is the frame rate of the detector.

Objects of the present disclosure will become apparent in light of thefollowing drawings and detailed description of the disclosure.

BRIEF DESCRIPTION OF THE DRAWINGS

For a fuller understanding of the nature and object of the presentdisclosure, reference is made to the accompanying drawings, wherein:

FIG. 1 is a schematic diagram of the basic hardware configuration foruse by the present disclosure;

FIG. 2 is a schematic diagram of the hardware configuration of FIG. 1with a three-layer detector structure wherein the beam selector with afixed focal point blocks primary x-rays from selected locations of therear detector;

FIG. 3 is a schematic diagram of the hardware configuration of FIG. 1with a three-layer detector structure wherein the beam selector with afixed focal point blocks scatter from selected locations of the reardetector;

FIG. 4 is a side view of a portion of a beam selector with an adjustablefocal point using compressible materials that blocks scatter fromselected locations of the rear detector;

FIG. 5 is a side view of a portion of a beam selector with an adjustablefocal point using x-ray-absorbent beads that block scatter from selectedlocations of the rear detector;

FIG. 6 is a side view of the portion of the beam selector of FIG. 4adjusted to a focal point;

FIG. 7 is a perspective view of a beam selector with a movable focalpoint using rigid materials and showing some of the tubes;

FIG. 8 is a top view of the beam selector of FIG. 7 ;

FIG. 9 is a side view of a beam selector of FIG. 7 adjusted to a nearer,center focal point;

FIG. 10 is a side view of a beam selector of FIG. 7 adjusted to amiddle-distance, center focal point;

FIG. 11 is a side view of a beam selector of FIG. 7 adjusted to afarther, center focal point;

FIG. 12 is a side view of a beam selector of FIG. 7 adjusted to amiddle-distance, left edge focal point;

FIG. 13 is a side view of a beam selector of FIG. 7 adjusted to amiddle-distance, right edge focal point;

FIG. 14 is a side view of a beam selector with a movable focal pointusing stacked plates;

FIG. 15 is a side view of the beam selector of FIG. 14 adjusted to adifferent focal point;

FIG. 16 illustrate unknown regions

FIG. 17 illustrates an embodiment of resolving for the unknowns in the zaxis.

FIG. 18 is a side view of a detector assembly having a scatter-blockingbeam selector and rear detector cells only at the base of the holes;

FIG. 19 is a schematic view of the x-ray source using the total internalreflection of polycapillary tubes to produce x-ray beams emitted fromdifferent locations;

FIG. 20 is a schematic view of an x-ray source using magneticdeflection;

FIG. 21 is a schematic diagram of a system using crystal interferometry;

FIG. 22 is a schematic representation of voxels as assumed by thepresent disclosure, where D_(x)=D_(y)≠D_(z);

FIG. 23 illustrates a NanoTube field emitter system;

FIGS. 24A-C illustrate configurations of field emitter activation zones;

FIG. 25 is a flow diagram of the basic steps of the 3D imaging methods;

FIG. 26 is a flow diagram of the basic steps of the 3D imaging methodswith a step added to separate primary x-rays and scatter;

FIG. 27 is a flow diagram of the basic steps of the 3D imaging methodswith steps added for 2D and 3D functional imaging;

FIG. 28 is a flow diagram showing multidimensional imaging from at leasttwo 2D images taken at two different x-ray emitting positions;

FIG. 29 is a schematic of a steering electron beams in a x ray tube withprogrammable field emission units.

FIG. 30 illustrates a number of pixels along a projected path.

FIG. 31 illustrates a configuration of x ray microscopy.

FIG. 32 is an example of an x-ray source.

DETAILED DESCRIPTION

A generalized spectral imaging description is described in Handbook ofMedical Imaging by J. T. Dobbins III, Image Quality Metrics for digitalsystems, 2000, pp 161-219, a projected signal from a detector element,at energy level Ω

∈{E1,E2,E3 . . . }

D(Ω)=qΩ∫ΦΩ(E)exp[−μ1(E)t1−μ2(E)t2−μ3(E)t3 . . . ]×SΩ(E)dE,  (1)

where q is the number of incident photons, Φ is the normalized incidentenergy spectrum, and S is the detector response function. Linearattenuation coefficients and integrated thicknesses for a number ofmaterials that make up the object are denoted μ and t, which attenuatethe X ray beam according to Lambert-Beers law. If we define attenuationof the subject along a projected path in a measurement unit μ(E)t wheret=time of a measurement unit, then when the measurement unit is a pixelpitch, μ(E)t is then the attenuation from a pixel, named as X. If thereare p pixels along the projected path measured by a detector element,the equation can be rewritten as

D(Ω)=qΩ∫ΦΩ(E)exp[−X ₁ −X ₂ −X ₃ . . . −X _(z) ]×SΩ(E)dE,  (2)

For example, as illustrated in FIG. 30 , z=12.

The basic method is to solve linear equation system(s) with m×n×pvariables, and m×n×p equations. Current CT methods assume that the pixelhas a size Xa=Xb=Xc, we will extend to the case where Xc is not equal toXc. When looking at the x ray passing the pixels, current methods, takethe value 1 or 0, 1 for the ray passing through the pixel, 0 if not.When passing through the volume will always the same. In the presentdisclosure, before conducting acquisition, a registration is made ateach angle and x ray emitting position. (I,j) receives signal passingthrough 1, 2, . . . 12, each subpixel transmission can be calculated.Assuming that inside each pixel, the transmission is uniform andproportion to the volume.

The present disclosure includes methods and systems for producingthree-dimensional (3D) images of a subject using two-dimensional (2D)x-ray detectors, by removing scatter, by a calibration method and a new3D reconstruction method. The methods and systems can include executiona number of steps, some of which are optional, as described below: (1)calibration, (2) acquiring 2D images from at least two different x-raysource locations relative to the subject, (3) processing the 2D imagesto separate primary x-rays from scatter, (4) functional processing ofthe 2D images, (5) processing to produce 3D images from the 2D images,functional processing of the 3D images, and/or presentation of theacquired information.

The preferred basic hardware employed by the present disclosure is shownschematically in FIG. 1 . It includes an x-ray source 12 and atwo-dimensional (2D) x-ray detector assembly 14. The image subject 2 ispositioned between the source 12 and detector assembly 14. The detectorassembly 14 communicates with a processor 15, including memory 19. Theprocessor communicates with user input and outputs 17, including, forexample, a display screen, keyboard and or mouse as would be understoodby a person of skill in the art. Other user interface elements can alsobe used as would be understood by a person of skill in the art from thepresent disclosure.

As mentioned in step (2) above, a number of 2D images are acquired. Foreach image, the location 16 from which the x-rays 30 are emittedrelative to the subject 2 is moved in a plane 18 parallel to the planeof the detector assembly 14. Consequently, the x-ray source 12 includesmechanisms for such motion, as described in detail below. The locationfrom which the x-rays 30 are emitted is referred to as the emittinglocation 16 in the remainder of the present specification.

As mentioned in step (3) above, primary x-rays and x-ray scatter areseparated in the 2D images prior to using the images. Typically, scatteris removed and discarded, which is why the separation is generallyreferred to as scatter removal in the remainder of the presentspecification. In some cases, however, scatter is used separately in,for example, trans vascular imaging, material differentiation, andidentification and inspection for better visualization of low atomic ztissues.

Primary X-Ray and Scatter Separation Methods

With reference to FIG. 2 , when x-rays 30 from the source 12 impact onthe subject 2, a portion of the x-rays 30 passes through the subject 2directly to the detector assembly 14 without a change in the directionof propagation. These are called primary x-rays 32 and convey trueinformation about the attenuation properties of subject 2. The remainderof the x-rays 30 are randomly scattered as a result of interactions withthe material of the subject 2. These are called scatter 34 and distortthe true information. In some cases, such distortion is used for formingseparated images which can be useful in representing true informationabout the subject as well.

Under some circumstances, scatter can be ignored. For example, where thedetector assembly 14 is a single 2D detector 20 that will receive bothprimary x-rays 32 and scatter 34, it can be assumed that scatter 34 ispresent but in a sufficiently small amount that qualitatively correct,yet quantitatively inaccurate, imaging results can still be obtainedunder certain circumstances. To what extent the amount of scatter 34 isacceptable is case-dependent and must be determined by a case-specificanalysis.

Primary x-rays and x-ray scatter can also be separated in the timedomain where a fast x-ray source is used. This method employs thecharacteristic that primary x-rays 32 travel in a straight line from thesource 12 to the detector assembly 14, taking the least amount of timein transit. Because scatter 34 does not travel in a straight line fromthe source 12 to the detector assembly 14, it takes a longer time toreach the detector assembly 14. Consequently, x-rays reach any givendetector cell 28 continuously over a period of time, where only thefirst x-rays are the primary x-rays. All others are scatter.Unfortunately, x-ray generation that allows for this time of analysis isextremely expensive and not practical.

Apparatus and methods described in the Chao's disclosure in US patentnos. U.S. Pat. Nos. 6,052,433A, 5,648,997A, 5,771,269A, 6,173,034B1,6,134,297A may be used in scatter removal with any embodiment of thepresent disclosure including examples where x ray source emittingpositions are close together enough that the use of beam selecting meansare not affected essentially. The foregoing disclosures are incorporatedherein by reference in their entirety.

In one example of this method, the source 12 is capable of generatingx-rays in extremely short pulses, for example, on the order of apicosecond in duration, and the detector assembly 14 is a 2D detector 20capable of extremely fast image capture, on the order of a picosecond.The captured image includes at least the primary x-rays 32 and thescatter 34 that reaches the detector 20 during the capture time window.If the capture window is short enough, the amount of scatter 34 in thecaptured image is minimized. As the capture window becomes shorter,scatter 34 becomes a smaller component of the captured image.

Another method is separation of primary x-rays 32 and scatter 34 in thefrequency domain. It employs the characteristic of some materialswherein the x-rays are modulated in space when passing through thematerial by attenuation. The x-rays 30 from the x-ray source 12 aremodulated with a different, typically much higher, frequency than thatof scattered x-rays from the subject 2. The high-frequencyattenuation-modulated x-rays, which are the primary x-rays, and thescattered x-rays, can be separated by post-image processing.

The present disclosure contemplates that any form of x-ray modulator canbe used. One example is a high-spatial-frequency pattern boardconsisting of attenuating blockers.

In another example, to separate x-ray signals by primary x-rays andscatter, in the spatial, frequency, or time domain, the modulator is aphase retarder, diffractive grating beam splitter withhigh-spatial-frequency pattern. In one form, such a modulator iscomprised of crystals or x-ray optics such as beam splitters, forexample, a kinoform structure. Additionally, such materials can bemodulated or tunable by acoustic waves in at least one dimension. Inanother example, the modulator is a MEM device, which operates as ahigh-spatial-frequency pattern board consisting of attenuating blockers,phase modulators, or polarity modulators. Such modulator can operate ina static or tunable state depending on the application.

Another example is that an interferogram is generated by applyingmodulation in space, time, or frequency, or combinations of two or moredomains, so that primary x-rays and scatter are separated based on thedifference in frequency, or time or spatial location.

In another mechanism, x-ray beams from two or more x-ray sources arecombined using a transmissive beam combiner. The interference fringes ofx-ray sources at an energy level or at a specific wavelength of x-raymoves as the phase difference between x-ray beams of the x-ray sourceschange relative to each other or alternatively as the wavelength of thex-ray changes. The result is the modulated interferogram creating amodulated spatial pattern, which can be separated from that of scatteredx-rays.

Similarly, one or more gratings are used to produce two or moreminisources of the x-ray source 12. The grating can be a dynamic ortunable grating, such as those formed as MEMs, or they can be static,such as a crystal grating. X-ray beams from the minisources interferewith each other and as the grating is tuned, a moving interferencefringe pattern is created. The result is the modulated interferogramcreates a modulated spatial pattern.

Similarly, a grating or a beam splitter can be used to split at leastone x-ray source 12 into two or more minisources with variable phases.As the wavelength or energy of the x-ray changes, the interferencepattern moves in space. The scatter and the primary x-rays are separatedas the scattered x-ray are separated from the signals which result inthe interference pattern measurements.

Another method of scatter and primary x-ray separation is described indetail in U.S. Pat. No. 6,134,297. The detector assembly 14 is athree-layer structure of a front 2D detector 22 closest to the source12, a 2D beam selector 24, and a rear 2D detector 26, as shown in FIG. 2. The combination of primary x-rays 32 and scatter 34 reach and passthrough the front detector 22. The beam selector 24 allows only scatter34 through to selected locations 54 of the rear detector 26.

The beam selector 24 can be an array of cylindrical shapes 62 comprisesx-ray-absorbent material and supported by a thin plastic sheet 60 havingnegligible x-ray absorption. The cylinders 62 are fabricated such thattheir axes are aligned with the travel direction of the primary x-rays32. As a result, the cylinders 62, within their cross-sectional areas,block all x-rays coming directly from the x-ray source 12. Thus, eachcylinder 62 produces a selected “shadowed” location 54 on the rear x-raydetector 26 where the strength of the primary x-rays 32 is essentiallyzero, while the strength of the scatter 34 is essentially unaffected.

Because the cylinders 62 have a finite size, a small portion of scatter34 will not reach the shadowed locations 54. However, as long as thecylinders 62 are small, this scatter 34 can be controlled to benegligibly small. If the cylinders 62 are too large or there are toomany, too much scatter 34 would be prevented from reaching the reardetector 26. The more cylinders 62 there are in the beam selector 24,the greater the accuracy of the measurement at the rear detector 26.

The cylinders 62 are fabricated such that their axes are aligned withthe direction of the travel of the primary x-rays 32, which means thatthe cylinders 62 are not parallel to each other, but are radial to thex-ray source 12.

The material of the cylinder 62 must ensure that substantially allprimary x-rays 32 are absorbed and, further, that it does not produceany secondary x-ray emission or cause any additional scattering. To meetthese requirements, chemical elements with a medium atomic number Z arepreferred, for example, materials with Z between 20 and 34. Thecylinders 62 can also have a multilayer structure, with a high-Zmaterial in the core and a medium-Z material outside. The high-Zmaterial absorbs x-rays most efficiently and any secondary x-rayemissions from the core material are efficiently absorbed by the outsidelayer without inducing further secondary emissions.

The thickness or the height of the cylinders 62 is dependent upon thex-ray energy, where higher energy requires thicker cylinders. In lowerenergy x-ray imaging, for example, in mammography, the cylinders 62 canbe thin disks.

The above-described detector assembly 14 is used to remove scatter 34from the image as follows. A low-resolution scatter image is read fromthe selected locations 54 of the rear detector 26. A low-resolutioncomposite image is read from chosen locations 56 of the rear detector 26that receive both primary x-rays 32 and scatter 34 and that uniformlycover the entire image plane of the rear detector 26 and are close tothe selected locations 54. The scatter-only image is extended to thechosen locations 56 by interpolation. The interpolation does not causesignificant error because of the physical nature of the scatter 34. Aslong as there are a sufficiently large number of data points, the errorincurred due to interpolation is negligible in comparison with othererror sources, such as statistical fluctuations of x-ray photon numbers.

The scatter-only interpolated image is subtracted from thelow-resolution rear composite image to produce a low-resolution primaryx-ray image at the chosen locations 56. A low-resolution primary x-rayfront detector image is calculated from the front detector locations 40aligned with the chosen rear detector locations 56. A low-resolutionscatter image is determined by subtracting the low-resolution primaryx-ray rear detector image from the low-resolution front detectorcomposite image. A high-resolution scatter image is calculated byinterpolating the low-resolution scatter image. The high-resolutionscatter image is subtracted from the high-resolution composite image toproduce a high-resolution primary x-ray image.

The above-described beam selector 24 must remain fixed relative to theemitting location 16 such as illustrated in FIG. 1 because the cylinders62 must remain aligned with the source x-rays 30 from the emittinglocation 16 to work properly. The detectors 22, 26 do not have to move,but will need to move if the beam selector 24 is fixed to the detectors22, 26. If the emitting location 16 moves a small enough distance thatthe cylinders 62 are sufficiently aligned with the source x-rays 30,then the beam selector 24 does not have to move.

Another method of scatter removal is described in detail in U.S. Pat.Nos. 5,648,997 and 5,771,269. The detector assembly 14 is a three-layerstructure of a front 2D detector 22 closest to the source 12, a 2D beamselector 24, and a rear 2D detector 26, as shown in FIG. 3 . Thecombination of primary x-rays 32 and scatter 34 reach and pass throughthe front detector 22. The beam selector 24 allows only primary x-rays32 through to selected locations 38 of the rear detector 26.

In the simplest configuration, the beam selector 24 is a sheet 46 ofx-ray-absorbent material having a large number of straight through-holes48. The holes 48 are fabricated such that their axes are aligned withthe travel direction of the primary x-rays 32, which means that, becausethe x-rays are emitted from essentially a point source, the holes 48 arenot parallel to each other, but are radially aligned to the x-ray source12.

Because of this alignment, the holes 48 permit all x-rays travelingalong the axes of the holes 48 to pass through, while almost all x-raystraveling in directions deviating slightly from the hole axes arecompletely absorbed by the bulk material of the beam selector 24. Thus,only the primary x-rays 32 reach the rear detector 26. Because the holes48 will always have a finite size, a small portion of scatter 34 willreach the rear detector 26. However, as long as the hole size 48 issmall and the thickness of the beam selector 24 is sufficiently large,this portion of scatter 34 can be controlled to be negligibly small incomparison with other sources of error.

Preferably, the holes 48 are as small as practical. If the holes 48 aretoo large, they will not prevent enough of the scatter 34 from reachingthe rear detector 26. Preferably, there are as many holes as practicalin the beam selector 24. The more holes 48 there are, the greater theaccuracy of the measurement at the rear detector 26.

The material of the beam selector 24 must ensure that all scatter 34 isabsorbed and that, except for the primary x-rays 32 passing through theholes 48, none of the other radiations, including scatter 34 andsecondary emissions caused either by primary x-rays 32 or by scatter 34,reach the rear detector 26.

The above-described detector assembly 14 is used to remove scatter 34from the image as follows. A low-resolution primary x-ray image is readfrom the selected locations 38 of the rear detector 26. Ahigh-resolution composite (primary x-rays 32 and scatter 34) image isread from the front detector 22. A low-resolution front detectorcomposite image is either read from or calculated from the frontdetector locations 40 aligned with the selected rear detector locations38. A low-resolution front detector scatter image is determined bysubtracting the low-resolution rear detector primary x-ray image fromthe low-resolution front detector composite image. A high-resolutionfront detector scatter image is calculated by interpolating thelow-resolution front detector scatter image. The high-resolution frontdetector scatter image is subtracted from the high-resolution frontdetector composite image to produce a high-resolution primary x-rayimage.

Because only the selected locations 38 on the rear detector 26 are used,an alternative structure for the rear detector 26 is to place one ormore detector cells at the base of each hole 48 rather than using anentire 2D detector with most of it unused.

The above-described beam selector 24 must remain fixed relative to theemitting location 16 because the holes 48 must remain aligned with thesource x-rays 30 from the emitting location 16 to work properly. Thedetectors 22, 26 do not have to move, but will need to move if the beamselector 24 is fixed to the detectors 22, 26. If the emitting location16 moves a small enough distance that the holes 48 are sufficientlyaligned with the source x-rays 30, then the beam selector 24 does nothave to move.

If there are only a very small number of emitting locations 16, the beamselector 24 can have a set of holes 48 with a different focal point foreach of the emitting locations 16. The holes 48 for the different setshave different selected locations 38 so that scatter 34 resulting fromone source location does not affect the signal from the other sourcelocations. This configuration of beam selector 24 does not have to movewith the emitting location 16.

The following beam selectors 24 have adjustable focal points so that thebeam selector 24 does not have to move as the emitting location 16moves.

The beam selector 24 of FIG. 4 comprises x-ray-absorbent material(s)with a large number of straight through-holes 70. Each hole 70 is formedby a rigid, hollow tube 76 made of one or more x-ray-absorbent materialsin layers 78 with different x-ray-energy absorbing characteristics. InFIG. 4 , the tubes 76 can comprise three layers 78 a, 78 b, 78 c ofdifferent x-ray-absorbent materials. Multiple x-ray-absorbent materialsprovide efficient absorption for multiple-energy systems or forsingle-energy systems where the energy is spread over a wide range, forexample 15 KeV to 500 KeV. Possible materials include tungsten (W), tin(Sn), and copper (Cu). When the materials are selected to absorb mostefficiently at different energy levels, the thickness of the tube wall78 can be minimized.

The region 80 between the tubes 76 comprises a compressible,x-ray-absorbent material, for example, an elastomer mixed with anx-ray-absorbent material. As with the tube walls 78, the region 80 cancomprise multiple layers with different x-ray-energy absorbingcharacteristics. In FIG. 4 , the region 80 comprises three layers 80 a,80 b, 80 c of different x-ray-absorbent materials, for example, anelastomer mixed with tungsten 80 a, an elastomer mixed with tin 80 b,and an elastomer mixed with copper 80 c. The ratio of elastomer tox-ray-absorbent material must be chosen such that the compressioncharacteristics of the elastomer and the absorbing ability of thex-ray-absorbent material are maintained. Typically, the ratio ofelastomer to x-ray-absorbent material is in the range of from 1:1 to1:4.

Alternatively, as shown in FIG. 5 , the region 80 between the tubes 76is blocked at both sides by plates 92 of polymers, elastomer, smartmetals, and/or some combination of these. The region 80 is filled withbeads 90 or powders of x-ray-absorbent materials. The thickness of beads90 or powders is adjusted as necessary to attenuate x-rays as desired.

Alternatively, smart metals and/or smart wires connect the tubes 64. Thesmart metals and/or wires can stretch, compress, or bend in at least onedimension under computer control.

Typically, in its quiescent state, the holes 70 are parallel to eachother and have no focal point. In order to create a focal point, the topedges 86 of the four sides 84 are displaced inwardly, as in FIG. 6 .Assuming that the four sides 84 are of equal length and that thecompressibility of the region 80 is constant throughout the beamselector 24, displacing all four edges 86 by the same amount willproduce a focal point above the center of the beam selector 24. Thedistance of the focal point will depend on the amount of displacement.By varying the amount of displacement of each edge 86 individually, thelocation of the focal point can be changed in three dimensions. Toachieve a focal point (as opposed to a focal line), at least twoadjacent edges 86 must be displaced.

Alternatively, the region 80 between the tubes 76 is notx-ray-absorbent. As long as the tube walls 78 comprised of one or morex-ray-absorbent materials block substantially all of the scatter 34 fromthe selected locations 38, the region 80 only needs to be filled with apolymer, elastomer, and/or any materials which can provide enoughstructural support.

Another example of a beam selector is a focused or parallel grid, linegrid, or crisscross grid formed by two line grids. A line grid isgenerally comprised of alternating lines of x-ray-absorbing material andspace, or materials of non-x-ray-absorbing properties andlow-x-ray-absorbing properties. The x-ray-absorbing material lines blockthe primary x-rays and the alternate spaces allow the transmission ofprimary x-rays. A crisscross pattern or line grid version may bevariations of the 2D beam selector mentioned above, where the layoutwhere the primary x-ray has come through is determined by the area wherelow-x-ray-absorbing materials or non-x-ray-absorbing materials areplaced. A focused grid is where the regions of the grid where theprimary x-rays pass through is designed so that the geometry of x-rayfan beam is taken into consideration. The focal point of the primaryx-ray regions is aligned with the x-ray fan beam passage way, so thatprimary x-rays are transmitted at distributed regions on the detector.The configurations described here are slight variations of previouslydescribed beam selectors using commonly used grids in x-ray collimators.

Another embodiment of a beam selector 24 with an adjustable focal pointis shown in FIGS. 7-13 . The beam selector 24 allows only primary x-rays32 through to selected locations 38 of the rear detector 26.

As shown in FIG. 7 , in one configuration, the beam selector 24comprises two or more rigid, x-ray-absorbent plates 100, 102 with alarge number of rigid, hollow tubes 104. The tubes 104 are formed ofx-ray-absorbent materials or polycapillary tubes capable of totalinternal reflection of beams at the restricted critical angle ornon-x-ray absorbent materials that do not generate scatter so long thatsuch setup does not allow scatter to pass through the tubes 104 andupper plate 102 and impinge on the detector at the bottom of othertubes. The tubes 104 can be formed like those of the beam selectorembodiment of FIG. 4 .

The beam selector embodiments above that allow only primary x-rays 32through to selected locations 38 of the rear detector 26 employ hollowtubes 104 with x-ray-absorbent walls. The present disclosurecontemplates that alternatives to the tubes 104 can be used.

Optionally, for the embodiments shown in FIG. 7 , beads or powders ofx-ray-absorbent materials are added in a pocket made of metal or polymersurrounding each tube 104 immediately above the rear detector 26. Thethickness and width of the pocket are designed to block substantiallyall (99.99%) primary x-rays 32 and scatter 34 from reaching the reardetector 26 around the tube 104, thereby creating an area devoid of anyx-ray signals surrounding the selected locations 38.

Optionally, for the embodiments employing elastomers, beads and/orpowders of x-ray-absorbent materials can be used between the plates 100,102 without elastomers.

In one configuration of an adjustable focal point beam selector, thebase plate 100 has a square matrix of apertures 106 of a given pitch andthe focusing plate 102 has a square matrix of apertures 108 with a pitchthat is slightly smaller than that of the base plate 100, as seen inFIG. 8 . For example, if the base plate pitch is 10 mm, the focusingplate pitch can be 9.8 mm.

The base plate apertures 106 are slightly larger than the diameter ofthe tubes 104 so that the tubes 104 can tilt within the aperture 106 toan angle of about 11° from vertical for a focal length of 500 mm. Thefocusing plate apertures 108 are slightly larger than the diameter ofthe tubes 104 so that the tubes can slide within the apertures 108 andtilt within the aperture 108 to an angle of up to about 11° fromvertical for a focal length of 500 mm.

When the focusing plate 102 is centered on the base plate 100 such thatthe center tube 104 is vertical, the focal point is directly above thecenter of the focusing plate 102. As shown in FIGS. 9-11 , moving thefocusing plate 102 vertically changes the vertical position of the focalpoint. As shown in FIGS. 13 and 13 , moving the focusing plate 102horizontally changes the horizontal position of the focal point.

In one configuration, the beam selector 24 has an array of totalinternal reflection elements, such as polycapillary conduits, where onlyx-ray beams that strike the conduits at a very narrow range of criticalangles are passed through to the rear detector 26.

Another example of a beam selector 24 with an adjustable focal point isshown in FIGS. 14 and 15 . In this embodiment, the tubes are eliminatedand perforations 116 are located in multiple stacked plates 114 ofx-ray-absorbent materials with thicknesses designed to absorbsubstantially all of x-rays. The perforations 116 align to form theholes 118 that pass primary x-rays 32 to the rear detector 26. Theplates 114 are moved along each other by 2D actuators in order to shiftthe focal point.

In another configuration, the beam selector 24 comprises x-ray optics,for example, an array of diffractive elements such as static andspatially-tunable or modulated MEMS or deformable MEMS, and crystals,such as ultrasound modulated spatially-tunable crystals for x-raytransmission at selected regions, or diffractive gratings and relatedx-ray optics.

When the direction of the incoming x-ray beam illuminating the crystalis varied, the modulation may change or remain the same correspondinglyto allow primary x-rays to pass through.

The beam selector 24 can have an array of refractive elements, such asx-ray optics or x-ray lens assemblies.

Separating primary x-rays or scatter can be based on the critical anglesfor incident x-ray beams. One bean may be within the critical angle andthe other beam may be outside of the critical angle combined with theuse of x-ray optics. For example, using refractive x-ray lenses as partof the beam selector, separation of primary x-ray or scatter may bebased on the acceptance angles for incident x-ray beams. In someinstances, such a beam selector could be part of a front window of therear detector which allows x-rays of selected regions to pass throughand be detected, thereby separating primary x-rays and scatter.Alternatively, the direction of the signals received at the reardetector corresponds to the incident angle of the x-ray beams, and witha known x-ray emitting location, geometry, and predicated incoming x-raybeam angle, the primary x-rays and scatter are differentiated at therear detector.

The configurations described above that employ scatter blocking beamselectors can also use a full 2D rear detector 26. In theseconfigurations, the detector cells 28 that are not aligned with theholes 48 are not used because they are blocked from receiving any x-raysby the beam selector 24.

Alternatively, as shown in FIG. 18 , rather than employing a full 2Ddetector, the rear detector 28 comprises groups 42 of one or moredetector cells 28 where each group 42 is positioned at the base of abeam selector hole 48.

A variation of the configuration comprising the beam selector sandwichedbetween two detectors is to use a dual-detector layer. In one example,the front detector is selectively transmissive of either primary x-raysor scatter while being able to detect the projected image that includesboth primary x-rays and scatter. Areas of pixels which are transmissiveregions comprise at least one transmissive pixel. In one example, thefront detector has a checker board pattern, with alternatingtransmissive where primary x ray can pass through and opaque regions.The alternating pattern and dimensions of the two regions are determinedby the requirement of the application and what percentage scatter needsto be removed or separated from the primary x-rays. In one example,below a transmissive area which comprises one or more pixels, a singlediode or multiple diodes with one pixel in each diode may be used asdetectors. Alternatively, a small detector with one or multiple pixelsmay be used. Thus, in place of a planar rear detector, there need not bea single detector, but rather multiple small detectors, each smallerthan the front detector.

In another configuration, each pixel of the front detector is partlytransmissive and partly senor. The rear detector captures thetransmitted signal coming out of the front detector.

In another variation, such transmissive or transparent features may bebuilt into the rear detector. Therefore x-rays may be absorbed ortransmitted in the rear detector before the sensing region of the reardetector.

Alternatively, such a detector assembly is built into one detector withfeatures built-in as described above and dual detecting planes.

Another variation of a beam selector uses filters for reducing theproportion of scatter reaching the second detector layer and a firstsensor attached to the filter to capture both primary x-rays andscatter.

Yet another example of a beam selector 116 placed between the frontdetector and rear detector comprises two or more plates, 114, asillustrated in FIG. 14 and FIG. 15 . Each of the plates comprises anx-ray absorbing material, such as tungsten, copper, or an alloy. Thethickness of each plate may be such that either alone or combined, itcan absorb almost all of the x-rays at one or more energy levels. Theremay be a spacer between the plates to maintain a distance betweenplates. Optionally, an x-ray-absorbing frame around the four sides ofthe plates not facing the detectors is placed as a beam stopper.

Each of the plates may have holes of various shapes. If the first plate114 is the plate closest to the front detector, the second 114, third114, fourth 114, fifth plate 114, may have corresponding holes ofsimilar or larger size to those of the first plate 114.

In some instances, each plate 114 may be moved by one or more linearactuators to align the x-ray source and holes distributed across eachplate to allow the primary x-rays to pass through.

Alternatively, if the holes are larger on the plates closer to the reardetector, there may be no need to move the plates. Only areas on thedetector where the primary x-ray beam paths are projected are used asthe selected primary signals. Any x-ray signal outside of the beam pathwill not be considered in derivation of a high-resolution primary x-raysignal on the front detector. These areas where the primary x-ray beampath project may be predetermined by the relative location of the x-raysource to the detector and the position of the beam selector plates, orby measurements done without the imaged subject, but with varied x-rayemitting positions corresponding to imaging configurations with theimaged subject.

X-Ray Source and Emitting Locations

The present disclosure uses an x-ray source 12 that is capable ofemitting pulsed or continuous x-rays with controllable energies. Such anx-ray source can be conventional x-ray tube x ray sources, cold-cathodex-ray sources, metal-liquid jets, laser Compton x-rays, linearaccelerator based x ray source, or synchrotron or synchrotron-like x-raysources, coherent or partially coherent x-ray sources, monochromaticsources, broad spectrum x-ray sources and/or field emitter nanostructurebased x ray sources.

Examples of an x ray source include a plurality of individuallyprogrammable electron emitting units for emitting an electron beam whenan electric field is applied, an anode target that emits an x-ray beamwhen subjected to collision of the electron beams emitted from an x-raysource having a collimator.

Each electron emission unit can comprise an electron field emissionmaterial. Each electron field emission material can comprise ananostructured material. The electron field emission material cancomprise a plurality of nanotubes or nanowires. Examples of thenanotubes (including carbon, boron and/or nitrogen) comprise at leastone field emission material selected from the group consisting of sulfurand/or tungsten. Examples of the nanowires (including silicon,germanium, carbon, oxygen, indium, cadmium, gallium, Zinc, oxides and/ornitrides) can comprise at least one field emission material selectedfrom the group consisting of silicides and/or borides.

Examples of a Portable Field Emitter Nanotube Based x Ray Source X-RaySource for Portability

The present disclosure provides an x ray source, an example of which isillustrated in FIG. 32 . It includes

-   -   An x ray generator with high efficiency, low weight and compact        size. For example, a vacuum-sealed field emission tube, s-1. S-1        can comprise a cold cathode x ray tube;    -   Electric energy storage, for example, a condenser or capacitor,        for storing the electric energy of the pulse operation, s-3. S-3        can be built at a pre-stage, can work at a much lower voltage,        for example, between 5 kV and 10 kV.    -   A voltage Amplifier, S-2 for example, a high voltage pulse        transformer, to provide high energy pulse, for example, from 100        kV-150 kV.

The pulse duration may be controlled, for example, between 0.1 ms and 10ms. The corresponding current flowing in the tube may be between, 10 m Aand 1 A to provide the desired amount of energy for a quality x-rayimaging. The pulse width and the current is determined by the parametersof the capacitance of the condenser, the inductance of the pulsetransformer, and the V-1 characteristic curves of the tube.

Flash x-ray sources can be based on use of a field emission tube drivenby for example, a high voltage source which provides high currentoutput, to trigger the pulse transformer. For example, HV semiconductordevices, HV tryrister-triggered pulse transformer or integrated gatebipolar transistor, such as illustrated in FIG. 31 . Vacuum or Gaseoustype of devices, such as Thyratron Spark Gaps, are examples of HighVoltage (HV) devices, which can provide HV pulser with high currentoutput.

High Voltage Thyrister or IGBT can serve as the switching device, S-5.Two or more such HV devices may connect in series to reach the initialHV voltage required to provide high current output such as a voltage inbetween 5-15 KV.

For variable energy generation at the output of the x ray source,different targets on a rotating anode, each generates x ray of a certainenergy level, is used.

X ray tubes maybe based on cold cathode x ray tubes. For example, fieldemission x ray source, comprising a plurality of programmable electronemitting unit, where each electron emitting unit maybe focused by eithera cone shipped anode target or a focusing means such as a electrostaticlens, electrooptic focusing lens as described above. The gate meshillustrated in FIG. 23-24C is optional.

The present disclosure provides an x ray source comprising majorcomponents with typical parameters: including for example, a 2.5 kV DCpower supply, a high voltage current storage means such as a condenserwith a capacitance, for example, 2 μF, an electronic triggering circuit,a high voltage pulse transformer, in one instance, encapsulated inSylgard Silicone, and an x-ray tube, which may be a cold cathode x raysource. The total energy stored in the condenser for generating a singlex-ray pulse maybe 30 J when HV=50 kV. The dimension of the whole x-raysource maybe less or about 8″×8″×16″, with a weight of about <30 lb. Asmaller dimension is possible depending on the choice of each component.The X ray energy level can be between 125 KV to 250 KV.

Such an x ray source maybe used in a two dimensional imaging system, andmay have portable capabilities.

To provide a tomography system with no moving parts, as illustrated byFIG. 20 , the electron beams or charged particle beams maybe steered bymagnetic means such as magnetic plates or magnetic based steering deviceor metal means, electrooptic lens for example, in a system with multipledimensional imaging capabilities as described in the present disclosure.

The x ray source device may generate single or multiple energies. It canbe used for one or more of the systems described below,

-   -   A x ray system of single, dual or multiple energy    -   A x ray system with scatter and primary x ray separated        involving hardware and software    -   A x ray multiple dimension or 3D imaging system    -   x ray imaging system for diagnostics, industrial and research        applications.

Typical Parameters of the source

Focus: 0.1 mm to 10 mm

Screen area: 60 cm×60 cm

Typical scan point: 30×30 to 50×50, maximum

Dwelling Time at one spot: 0.01 ms to 10 ms

Typical Time at one spot: 1 ms

For one complete set of image, determined by the detector frame rate.

The source 12 can emit two or more consecutive x-ray pulses withcontrollable energies for each imaging operation: a high-energy pulse atan average energy level H followed by a low-energy pulse at an averageenergy level L. Each pulse has a single, reproducible energy spectrum,which comprises bremsstrahlung radiation and discrete line emissions.

In another configuration, the source 12 emits three consecutive pulsesfor each imaging operation: a high-energy pulse at an average energylevel H, followed by a medium-energy pulse at an average energy level M,followed by a low-energy pulse at an average energy level L. Each pulsehas a single, essentially unchanged energy spectrum.

In another configuration, the source 12 emits a pulse which has two ormore energy peaks each separated from the other energy peaks temporallywithin the pulse duration.

In another configuration, the source 12 emits a broad spectrum pulse.

Each of the detectors or x ray measurement assembly including thedetectors may be energy sensitive, such as photo diode arrays or energysensitive detectors or x ray spectrometer.

Movement of the x-Ray Emitting Locations

The x-ray emitting position may be moved by mechanical, electricaland/or magnetic energy means.

With reference to FIGS. 1 and 19 , the x-ray emitting location 16 canmove relative to the subject 2 in a plane 18 parallel to the detectorassembly 14. Alternatively, the present disclosure contemplates that theemitting location 16 can move relative to the subject 2 in a planeperpendicular to the detector assembly 14 or in a plane other thanparallel and perpendicular to the detector assembly 14. A mechanical,electrical or magnetic mechanism provides the desired motion.

In another configuration, the mechanism moves the emitting location 16either angularly, linearly, or a combination of both. The movement ispreferably done to solve the unknown pixels in the third dimension inthe region of interest 4 within the subject 2, preferably in integermultiples of pixel pitch, while minimizing the introduction of newunknown pixels in each movement and minimizing introduction of a totalnumber of new unknown pixels for the complete derivation of unknownpixels in the third dimension for the region of interest 4.

To minimize total imaging time and radiation exposure, the presentdisclosure contemplates, though does not require, that the mechanismprovides the emitting location 16 motion rapidly. For example, themechanism can provide this motion in increments of integer multiples ofpixel pitch (the distance between adjacent detector cells 28).

The present disclosure contemplates that the subject 2 is physicallymoved relative to the emitting location 16, particularly inapplications, such as industrial applications, where the subject 2 isalready in motion while being imaged. To minimize total imaging time andradiation exposure, each movement, either angular or linear, resolvesthe unknown pixels in the third dimension, preferably in integermultiples of pixel pitch and at the same time, new unknown pixels areintroduced in the process, the number of measurements required toresolve such unknown pixels in addition to the region of interest willneed to be taken into account.

One example of an x-ray source and movement system is a metal, graphene,silicon or carbon nano-tube-based field emitter, which is driven by anapplied electric field, not by temperature, that is also known as a“cold cathode”. In a CNT, the current is exactly and instantaneouslycontrolled by an applied voltage.

A CNT emitter, c-3, can be used in the x-ray tube including two or morefield emitter nanotubes such as carbon nanotubes, arranged vertically ona conductive substrate. In one configuration, the width of carbonnanotubes is in the nanometer range. The tip of the carbon nanotube iswhere the emission occurs. In one example, the CNTs are arranged asillustrated in FIG. 23 . A grid or a gate mesh structure, c-2, connectedto an electrode inside the tube wall is placed a small distance abovethe tips of the CNT. A voltage gradient may be applied externallybetween the grid mesh and the substrate. The voltage generates a verystrong electric field at the tips of the CNTs. The electric fieldstrength concentrated at the CNT tips forces field emission ofnegatively charged electrons to occur at the tips of the CNTs. When theexposure is switched on, the electrons are emitted from the CNT tips andfly toward the grid mesh, c-2. The majority pass through to beaccelerated by the anode high voltage, generating x-rays when theelectrons impact the anode c-1. The electro-optic focus lens, c-4, whichcan be inside, can be used to dynamically adjust the focal spot size ofthe x-ray emission or to move the focal spot on the anode to differentlocations.

As illustrated in FIG. 24A, different regions of CNT emitters may beactivated, combined with using the electro-optic electrodes c-4 tocontrol the electron beams to adjust the size of the focal point of theelectrons c-5 on the anode.

To control the size of electrons on the focal point, other types ofelectrostatic lens or electromagnetic lens can be used, for example, anelectromagnetic coil, einzel lens, quadrapole lens, magnetic lens, ormultipole lenses.

An additional electromagnetic means for example, an electromagnetic coilc-6 may be used to steer the beam so that the x-ray beams may be emittedfrom various locations on the anode from the XY plane parallel to thedetector. In another embodiment, for electronic beam steering, a set ofmagnetic plates, or metal plates may be used.

In one configuration, CNT emitter element c-3 comprising regions c-7,each comprising one or more field emitter nanotube emitters asillustrated in FIGS. 24A, 24B and 24C, different implementations of thisconfigurations. Each region may be activated and deactivatedindependently from the others. Electron beams originating from one ormore field emitter nanotube emitters may be steered electronically byelectrooptical devices or magnetically by magnetic plates orelectromagnetically by electromagnetic coils c-5 so that its focal spoton the anode may shift its position in a programmable pattern. Theoverall x-ray emitting positions on the anode may be extended in area ordistance by using two or more such regions, for instance, so that thepattern of x-ray emitting positions may be continuous throughout thex-ray emitting area on the anode.

With specific reference to systems that use a beam selector, such asbeam selector 24. Some beam selector embodiments must remain fixedrelative to the emitting location 16 because they have fixed focalpoints, and other beam selector embodiments do not have to remain fixedrelative to the emitting location 16 because they have adjustable focalpoints.

Several different mechanisms for providing a moving emitting location 16and, optionally, a moving detector assembly 14 are described below.

In one mechanism, two or more x-ray sources 14 are positioned atdifferent locations in the plane 18 and emit pulses sequentially fromthose locations. The detector assembly 14 is fixed. For this mechanism,the beam selector 24 must have either multiple fixed focal points or anadjustable focal point.

In other mechanisms, a single x-ray source 12 emits x-ray pulsessequentially from different locations on the plane 18. The presentdisclosure contemplates several different configurations that canaccomplish this.

In one configuration, a two-dimensional actuator physically moves thex-ray source 12 and the x-ray detector assembly 14. Preferably, theactuator can move the x-ray source 12 and the x-ray detector assembly 14an integer multiple of one detector pixel pitch in the plane 18 at orfaster than the frame rate of the detector assembly 14. For thisconfiguration, the beam selector 24 can have a fixed focal point.

In another configuration, a two-dimensional actuator physically movesonly the x-ray source 12. Preferably, the actuator can move the x-raysource 12 in the plane parallel to the detector assembly 14 inincrements of the minimum integer multiple of pixel pitch, which leadsto a planar motion of one-pixel pitch for an x-ray beam measurementlocation relative to the previous position, at or faster than the framerate of the detector assembly 14. For this configuration, the beamselector 24 must align with the emitting location 16. In someconfigurations, the beam selector 24 may need to have an adjustablefocal point.

In another configuration, a two-dimensional actuator physically rotatesonly the x-ray source 12 so that the emitting location 16 moves in anarc. Preferably, the actuator can rotate the x-ray source 12 theequivalent of an angle along the arc to simulate a planar motion of onepixel pitch, or an integer multiple of one detector pixel pitch at orfaster than the frame rate of the detector assembly 14. For thisconfiguration, the beam selector 24 must align with the emittinglocation 16 and, in some configurations, may need to adjust its focalpoint.

In another configuration, shown in FIG. 19 , the x-ray source 12 is asingle x-ray source 202 and the x-ray beam is moved to differentemitting locations 16 using total internal reflection of, for example,polycapillary tubes 204.

In other mechanisms, methods are used to deflect the electron beamwithin the x-ray source 12 to hit a different location on the anode,thereby causing the x-ray beam to be emitted from a different emittinglocation 16.

In one configuration, shown in FIG. 20 , a changing magnetic fieldgenerated by, for example, a solenoid coil 212 attached to the housingof the x-ray tube 210, deflects the x-ray beam. The energized coil 212produces a magnetic field and an associated Lorentz force on theelectron beam in the x-ray tube 210, shifting the impact spot on theanode target 214 from which x-rays are emitted. The emitting location 16moves due to the displacement of the focal spot of the cone beam on theanode 214. The result is that the emitting location 16 moves from onelocation to another. Careful control of the coil 212 can producemovement in as small as a one-pixel pitch in one or two dimensions.

In another configuration, the electron beam is deflected as it passesthrough charged metal plates. The direction of deflection depends on thepolarity and amount of charge of the plates.

In another configuration, a light source such as a light-emitting diode(LED) or laser are used as the source to generate the electron beam,which is then amplified by a multiplier tube. A light deflector such asoptics or mirrors, and acoustic optical deflectors are used to deflectthe light. An ultrafast laser may be used to generate an ultravioletemitter that emits ultraviolet light, a photocathode operably coupled tothe ultraviolet light-emitting diode that emits electrons, an electronmultiplier operably coupled to the photocathode that multiplies incidentelectrons, and an anode operably coupled to the electron multiplier thatis configured to produce X-rays. The ultraviolet emitter may be steeredat different angles to control the output of the electron beam which, inturn, changes the direction or the location of the x-ray beam emittedfrom the anode.

In another configuration, irradiating arrays of metal components, suchas nanowires, with intense femtosecond laser pulses produceshigh-brightness picosecond X-ray pulses. The emitting location 16 can bemoved by using optical steering devices to change the impact location ofthe laser beam on the metal components.

Multiple photocathodes or a multiplier may be used to collect thesteered laser beam and output an electron beam. In some instances, theelectron beam may be steered quickly via electronic, magnetic, optical,and/or crystal plus ultrasound, or other means. One way to do this is touse a 2D array of collimators, the holes being integer multiples of apixel pitch apart, raster scan x-ray beam using different mechanismsincluding magnetic means.

In another configuration, the x-ray source comprises micron-scale metalx-ray emitters which can be modulated and switch on and off to controlthe emitting location.

In another configuration described in U.S. Patent Publication No.2010/0189223A1, the x-ray source 12 is a digitally-addressed flat panelof x-ray sources, where one or more emitters are located in each pixelon a 2D flat panel.

In another configuration, ultrasound can modulate an x-ray beam in spaceand time. For example, space-time modulation of an x-ray beam is done bytotal external reflection on a YZ-cut of a LiNbO/sub 3/crystal modulatedby surface acoustic waves. The x-ray diffraction is determined by theamplitude and wavelength of the surface acoustic waves. An x-ray beam isaimed at the crystal and the crystal dynamically steers the beamdepending on how the diffraction properties of the crystal are changedby the modulating acoustic wave. The output location of the x-ray fromthe crystal may be varied. For example, the x-ray beam may be modulatedto exit from crystal at various locations on an XY plane.

In another configuration, relativistic electron beams with currentmodulations at a nanometer scale and below may be used to generatecoherent x-rays. The current modulation is produced by diffractingrelativistic electrons in a perfect silicon crystal, accelerating thediffracted beam and imaging the crystal structure, then transferring theimage into the temporal dimension via emittance exchange. The modulationperiod can be tuned by adjusting electron optics after diffraction. Thetunable longitudinal modulation can have a period as short as a fewangstroms, enabling production of coherent hard x-rays from a devicebased on inverse Compton scattering with total length of a few meters.

The spatial position of electron beam where the electron beam strikesthe anode, may be modulated by steering the position of the laser beamfrom the laser pump, which gives rise to the varied position of x-rayemitting position. The present disclosure also contemplates the use ofcombined imaging systems, for example, interferogram, partially-coherentand coherent x-ray imaging systems. For example systems used in phasecontrast imaging methods, coherent x-ray imaging or interferogram basedsystems upstream of the detector are modified to achieve varied x-rayemitting locations or to generate varied x-ray sources. For example,source grating, where the x-ray is split into multiple sources upstreamof an x-ray interferometer, may be used in the generation of x-rayemitting positions. This may be done either by a tunable grid whereselected x-ray emitting locations are blocked temporally by x-rayabsorbing regions. Such a tunable grid system can be designed into thesource grating hardware or can be separate. Another embodiment is that,downstream from the source, the x-ray emitting locations are variedeither by using a blocking mechanism similar to that at the source or bymoving or modifying any hardware downstream from the source, anywherebetween the source and detector, to achieve the end purpose ofmeasurements of regions of interest using varied illumination paths ofthe region of interest at different times. For example, G1 and G2gratings may be modified.

Another embodiment of x-ray interferometer is crystal interferometer, orinterferometer based on crystal elements.

As shown in FIG. 21 , an example of crystal interferometry or x-rayinterferometry based on crystals, consists of three beam splitters 150,152, 154 in a Laue geometry aligned parallel to each other. The incidentbeam 30 from the x-ray source 12, which usually is collimated andfiltered by a monochromator (Bragg crystal) 156 before, is split at thefirst crystal 150 by Laue diffraction into two coherent beams, areference beam 160 which remains undisturbed and a beam 162 passingthrough the subject 2. The second crystal 152 acts as a transmissionmirror and causes the beams 160, 162 to converge towards each other. Thetwo beams 160, 162 meet at the plane of the third crystal 154, which issometimes called the analyzer crystal, and create an interferencepattern, the form of which depends on the optical path differencebetween the two beams 160, 162 caused by the subject 2. Thisinterference pattern is detected by an x-ray detector 14 behind theanalyzer crystal 154.

By moving one of the hardware elements, for example, the Bragg crystal156, the x-ray emitting location 16, or the subject 2, and recordingprojections from different illumination path using methods describedabove, the 3D distribution of the refractive index and thus tomographicimages of the subject can be retrieved.

In contrast to the methods below, with the crystal interferometer thephase itself is measured and not any spatial alteration of it. Toretrieve the phase shift out of the interference patterns, a techniquecalled phase-stepping or fringe scanning is used. A phase shifter (withthe shape of a wedge) is introduced into the reference beam. The phaseshifter creates straight interference fringes with regular intervals, socalled carrier fringes. When the subject is placed in the other beam,the carrier fringes are displaced. The phase shift caused by the subjectcorresponds to the displacement of the carrier fringes. Severalinterference patterns are recorded for different shifts of the referencebeam 160 and by analyzing them the phase information modulo a can beextracted. This ambiguity of the phase is called the phase wrappingeffect and can be removed by so-called “phase unwrapping techniques”.These techniques can be used when the signal-to-noise ratio of the imageis sufficiently high and phase variation is not too abrupt.

As an alternative to the fringe scanning method, the Fourier-transformmethod can be used to extract the phase shift information with only oneinterferogram, thus shortening the exposure time, but this has thedisadvantage of limiting the spatial resolution by the spacing of thecarrier fringes.

“Coherence-contrast x-ray imaging” is another embodiment ofinterferogram, instead of the phase shift the change of the degree ofcoherence caused by the subject is relevant for the contrast of theimage.

Alternatively the Laue crystals can be replaced by Bragg crystals, sothat the beam does not pass through the crystal but is reflected on thesurface.

In some cases, two crystals instead of one are used to enlarge the fieldof view.

It is to be noted that in most, if not all phase contrast measurementsand interferometry methods, the amplitude of the x-ray input beam orintensity of the x-ray input beam may be modulated for somemeasurements.

The crystal-based interferogram or analyzer where the location of thebeam splitter can be steered using modulation by ultrasound, or otherenergy or electronic methods. For example, a liquid crystal basedcrystal analyzer can be used to adjust the beam splitting locations togenerate different illumination paths.

The present disclosure can be implemented in a portable format. Forexample, handheld, carry on version. Another example of the presentdisclosure in a portable version is a foldable system which adapted forfield inspection or diagnostics or image guidance or materialcharacterization.

Additional Advanced Hardware

It is one aspect of the present disclosure to include the use of

-   -   a Micro Electronic Mirror (MEM) device for use as 1) a static or        dynamic or tunable diffraction grating, in interferometer, or        as 2) primary modulator for generate high spatial frequency        x-ray beams, or as 3) collimator by adjusting critical angle for        the incident x-ray beam.    -   a crystal or liquid crystal based device for use as 1) a static        or dynamic or tunable diffraction grating, in interferometer, or        as 2) primary modulator for generate high spatial frequency        x-ray beams to separate scatter and primary x-rays, or as 3)        collimator by adjusting critical angle for the incident x-ray        beam.

Methods

An element of the present disclosure includes methods of rapid,high-resolution 3D image reconstruction of a region of interest in asubject by moving the position of the x-ray source emitting locationrelative to the detector in smallest possible increments, for example,pixel increments to only a small number of positions in the area ofapproximately n2 in the XY plane to acquire the image data necessary forresolving n2 unknown pixels along the Z axis. Where n is the highestunit number in the x or y axis describing the x ray emitting location ina 2 D plane. This allows for a system that requires fewer 2D projectionimages, less motion requirements, or less motion movement, or acombination of these elements, which means less time needed and lessradiation exposure to reconstruct multidimensional images than in theprior art.

In an x ray measurement using a 2D detector, the unit measurement of nmaybe that of a pixel pitch of the detector.

In some cases, the relative position of x ray emitting locationsadjacent to each other maybe less than one pixel pitch, but stillproduce measurements for the project path which may be used to resolvethe unknown voxels in the z axis perpendicular to detector.

In some cases, such as in an x ray microscopy device, the resolution ofthe camera, therefore the pixel pitch, maybe much larger than themeasurement unit or the resolution achieved in the subject. For example,in the nanometer or sub nanometer range. As a result, the measurementunit of n or the distance between adjacent x ray emitting positionsmaybe in the range of resolution to be achieved in the subject, forexample, sub-nanometer or single digital nanometern, or as small as theachievable resolution allows. This means that instead of a measurementof a pixel pitch for the distances between adjacent x ray emittingpositions, the distance may be many times smaller than a pixel pitch ofthe light camera. In some in cases, as small as the smallest x raywavelengths, for example, 0.01 nm.

The present disclosure also describes methods where the subject movesrelative to the x ray source and detector, such as a part on an assemblyline. The x ray 3D measurements maybe done by calibrating andregistering the x ray emitting positions, the element of detector,relative to the location of the part, as the part moves through theassembly line. 3D image acquisition maybe achieved independent of theunknown volume introduced by the movement of the part. The movement ofthe part may also be used as part of measurements done to resolveunknown voxels along the z axis. the number and locations of the x rayemitting position can be determined based on a similar strategy. Thegoal is to minimize the total area of the movement and to minimize themovement distance so that the depth or the z axis of a new unknownvolume introduced in the process of resolving the unknown voxels of theregion of interest is minimized.

The methods include the following elements:

Determine the number of 2D images needed for reconstruction of thedesired 3D images.

For example, if the 2D detector array has m×n detector cells, then anynumber p of 2D images, where p<n and p<m, and p is the third axisunknown variable, >3. This results in p number of 2D images with m×npixels each.

The basic method is to solve a linear equation system with m X n X pvariables and m×n×p equations. Current methods do not assume that eachvoxel in the region of interest of the subject is cubic, that is, thesides of each voxel are the same length, D_(x)=D_(y)=D_(z). The currentmethods apply to the case where the side in D_(z) is not equal to D_(x)and D_(y), shown schematically in FIG. 22 .

When the location of the x-ray emission position 16 relative to thedetector is shifted one pixel pitch in the x or y direction parallel tothe detector, the new emitting position can generate x-ray beams toilluminate the region of interest with different projected paths thanthe previous position. This means that, in each projected path, therewill be voxels with different spatial position within m×n×p m,n,p) inthe projected path, which leads to different x-ray measurement positionon the detector.

When such positions (i,j) on the detectors are correlated to the x-rayemission position 16 which is the spot on source 12, illustrated in FIG.1 , FIG. 14 , and FIG. 15 , as the subject is placed in the illuminationpath, different locations on the detector measures attenuated signals xof the projected path signals within the volume of the region of theinterest, each measurements is correlated to the relative position ofthe x-ray source and the detector as determined by previous derivedgeometric information. where each projected path is comprised of Pvoxels, each containing its specific attenuation coefficient.

In a volume of m×n×p where p is the third axis thickness or variable, inthe region of interest. Varying different x-ray source emissionpositions p times, different 2D projection images are generated so thateach pixel on the detector measuring projected line signals from theregion of interest contain unique set of pixels with fixed spatialrelationship between the pixel elements. Each 2D image is comprising atleast m×n measurements, are measured p times, thereby creating m×n×pknown measurements. The dataset is then used to resolve unknown voxelsin the m×n×p space which are variables responsible for generating variedx-ray measurements at corresponding locations on the detector at variedx-ray emitting positions.

In cases, where the unknown pixels of the subject are embedded in thevolume of known pixels of the subject, the x-ray emitting positions maybe shifted by a fraction of the pixel pitch to resolve the unknownpixels by creating projection paths sampling at finer steps, so that theprojected paths include at least one unknown pixel and, but at the sametime do not introduce new unknown pixels.

In addition, in cases where the resolution requirement is not as high asthe detector pixel size, it is within the present disclosure to definethe resolution required and the corresponding measurement resolution.For example, if the resolution required is 500 μm and each pixel on thedetector has a pixel pitch of 100 μm, it is one aspect of the presentdisclosure to move the x-ray emission location at 500 μm. When the 3Dimage acquired at 500 μm resolution is obtained for the region ofinterest, further measurement at finer steps, for example 100 μm, may beused to resolve finer details in the selected areas within the region ofinterest.

Furthermore, in situations where the material composition is determinedand anatomical markers and dimensions are derived either usingmultiple-energy material decomposition or using lower resolution 3Dimaging method, or is simply given, or derived based on known facts,such information can be used to select one or more regions for furtherfiner resolution imaging.

The following general steps can be performed to determine a 3D image:(1) Calibration; (2) 2D imaging; (3) 2D image scatter removal; (4) 2Dfunctional imaging; (5) 3D/multidimensional image calculation; (6)3D/multidimensional functional imaging; and (7) 3D/synthesized2D/multidimensional image presentation Each step is optional and thesteps can be performed in a different order.

Calibration

Before performing an image acquisition of the subject, a preliminarygeometric relationships can be determined for each x-ray source locationand the detector without a subject. As illustrated in FIG. 30 , Asubpixel or pixel transmission can be calculated for each cell (i,j) onthe detector 22, where i,j denotes the pixel or region x,y position onthe detector 22 where it receives the x-ray signal passing throughvoxels on the projected path, 1, 2, 3, . . . 11, 12, etc. The pixelpitch is the distance between the x-ray source movement steps or thedistance between adjacent x-ray emitting positions. Xa, the pixel pitchof the detector, may be the same as Xc, which is the resolution or theunit of measurement in depth or in the axis perpendicular to thedetector. In some instances, Xa may be smaller than Xc, the resolutionrequired in the Z axis, Xc. X ray emission positions move in Xcincrements in order to achieve the required resolution.

The basic method is to solve linear equation systems with m×n×pvariables, and m×n×p equations. Current CT methods assume that the pixelhas a size Xa=Xb=Xc, the present disclosure however extends to the casewhere Xc is not equal to Xc. When looking at the x ray as it passes thepixels, current methods take the value 1 or 0, 1 for the ray passingthrough the pixel, 0 if not. In the present disclosure, beforeconducting acquisition of the region of interest, a registration is madeat each angle and x ray emitting position (i,j) receives a signalpassing through 1, 2, . . . 12, each subpixel transmission can becalculated. Assuming that inside each pixel, the transmission is uniformand proportion to the volume.

This geometric calculation can be done in advance and stored in thecomputer. Alternatively, a general formula can be derived for how towrite the equation system in numerical format.

Starting with the equation system m×n×p equations, with m×n×p variables,for each direction of projection x ray source or x ray emittingposition, there are m X n projection data points, recorded by the m X ndetector cells. Each equation has m×n×p variables, there are (m×n)×pequations, thus the linear equation system should be solvable by eitheriterative method or matrix method.

In one preferred embodiment, if a subject of certain dimensions is to beimaged, the thickness or dimension measurement in the Z direction forthe region of interest can be provided ahead of the time. This can bedone by

-   -   1. a user input, or    -   2. data derived from, for example, mechanical measurement tools        or    -   3. by 2D x ray measurements at multiple energies to determine        the thickness based on a database established using previous        measured data of the same or similar materials, or    -   4. by simulated data, combined with one or more 2D images at        single energy or dual or more energy    -   5. x ray measurements at two or more different emitting        locations, or    -   6. by one or more non-radiation sensor.

In addition, the location of region of interest, relative to the x raysource and detector maybe determined using the techniques as describedin 1-6.

For example, a component size of 2 cm in xyz dimensions labeled withcontrast agents, or a region of interest in the bone and its immediateregion for arthritis identification, a strained stress region of thebone or certain area of a semiconductor assembly line, an area ofinterest in a luggage for security inspection, a region of interest fordetailed inspection of a part in an assembly line.

Based on the resolution Xc as illustrated in FIG. 30 required, then thenumber and location of x-ray emitting positions are determined for 3Dimaging of the subject. Prior to illuminating the subject from differentx-ray emitting locations, calculations or calculation and measurementsmay be performed to correlate the relative x-ray source locations andpixel locations on the detector for their corresponding projected pathswithin the volume where the region of interest in the subject is placed.Therefore, each projected path in 3D dimensions of the region ofinterest is related to the x-ray source location and pixels or pixelregions on the detector, where such projected x-ray signal is measured.

After the 2D images are taken, solve for the equation system m×n×pequations with m×n×p variables. Each emitting location of the x-raysource produces an image of size m×n. And each element of the detectorreceives x ray transmitted through a predetermined projection path, someof the elements receives x ray transmitted after attenuation at theregion of interest in a 3D space. There are a total of p third axisvariables or unknown pixels that need to be resolved to provide thecomplete 3D image of the region of interest. For example the totalattenuation variable

-   -   Attenuation X_(total) along any given projected path with the        total number of pixel numbers being Z is

X _(total) =X ₁ +X ₂ . . . +X _(z)  (3)

As illustrated in FIG. 30 . Each X out of the

X ₁ ,X ₂ . . . X _(z)

Corresponds to a unique subpixel or pixel in the m×n×p volume in theregion of interest.

There is an issue of completeness, for example, when the unknown pixelsoutside of the region of interest are introduced into the measurementsas the relative position of the x-ray source and the detector moves.There are three situations. (a) When the region of interest of thesubject is well inside the imaging area, completeness of the 3D image isassumed because the linear equation system usually gives the rightsolution. (b) When the region of interest extends beyond the imagingarea in one dimension, there is an issue that will be discussed below.(c) When the region of interest extends beyond the imaging area in twodimensions, there is also an issue further discussed below.

Assume that region A is the region of the interest and region B isadjacent to region A. But for acquiring projection data of the region A,we must deal with the data in the region B. Region A is surrounded byregion B by conducting a two-step imaging process in which x-ray sourcesare moved differently, or a two-step scan if x-ray emitting positionsare moved differently due to the electron beam being steered to scananode target at various locations. This generates x-ray at differentlocations. Information in the region B can be accurately gained withoutfurther extending to a larger region. The method of the presentdisclosure allows for minimizing x-ray measurements required forreconstructing the multidimensional image. For example, using atwo-direction move, or raster scan in 2D area, acquire projection imagesalong the X direction and the Y direction, NX data points and NY datapoints, total projections on the 2D plane, is (NAX+NBX)(NAY+NBY) asillustrated in FIG. 16 .

-   -   Region A has NAX×NAY pixels    -   Region B has NBX×NBY pixels    -   Total (NAX+NBX)(NAY+NBY)pixels=NAX×NAY+NAX×NBY+NAY×NBX+NBX×NBY,        new unknown NAX×NBY+NAY×NBX

For example, the region of interest A has an X dimension of 20 cm, a Ydimension of 20 cm, and a Z dimension of 20 cm. If the pixel pitch ofthe x-ray detector is 200 μm, there would be 1000 data points. In otherwords, there would be 1000 unknown data points to be resolved.Therefore, when the x-ray from the source is scanned in the XY planeparallel to the detector, x-ray images are sampled by at least 1000different emitting positions. The distance between each of the adjacentemitting positions can be chosen to produce the minimum number ofunknowns and, at the same time, minimize the sampling time.

For example, steering the x ray emitting position along the x axis and yaxis at 1000 first positions within an area of 33×33, acquiringmeasurements at each position, and the distance between adjacentemitting positions is the pixel pitch of the detector, which is 200 μm,

However as the x-ray emitting position changes to illuminate region A,unknown pixels in region B, the area surrounding region A are introduced

The solution to the above problem is to make the scan step finer in asecond position, so that all the information in region B can beaccurately gained without further extending to an even larger region.

The scanning of x-ray source or the x-ray emitting position remains inthe scanned area in first position, except now, the x-ray source isscanned in finer steps which illuminate x ray on the region of interestand unknowns in Region B. For example x rays can be sampled a fractionof movement step size, or a fraction of pixel pitch, from a previousx-ray emitting position. If each emitting position is different from theprevious emitting positions in the step described earlier, however atthe same time, are only illuminating the unknown pixels which need to beresolved within Region A and Region B, additional measurements which maylead to formulation of new linear equations which reveal newrelationships between existing unknowns in the region of interest RegionA or subsets of unknowns of Region A and those newly introduced inRegion B outside of region of interest in the step described earlier.The sets of linear equations may be derived to give rise to the completesolution as illustrated by equation (3), where each of the pixel in theprojected path comprising a unique set of pixel or voxels in the regionof interest A and in some parts outside of region A. The newlyintroduced unknown regions will have a depth or measurement in the zaxis such as p=n² as in FIG. 17 and the total volume of new unknownsintroduced is 4 n×n², n² being the depth of the region of the interest,the unit of measurement is the movement resolution or movement steps.For example, if there are 100 unknowns along the z axis in region A, thetotal unknowns introduced newly is 4×10×100 if the x ray emittingpositions have moved within 10×10 area in the XY plane. If however the xray emitting position moves in a larger area, for example, 15×15, thenthe furthest area from the original position will be higher, as aresult, new unknowns introduced will be a lot more, the total unknown isnow 100+4×15=160. 15×15 more than needed.

When the depth is n² 1000, xy direction move is square root of 1000,which is approximately 32, newly introduced unknown is 4×32=130 and thetotal unknowns is 1130, or approximately 34 or in order to resolve the1000 unknowns in the z axis, ideally x ray emitting positions will be at1130 different first positions, each would be pixel pitch apart from itsmost adjacent spots, the x ray emitting positions is less than 34 orapproximately 32 pixel pitch away from the original position in the xdirection and the y direction in order to minimize the imaging time.

For the step of finer scan, second points, the x ray emitting positioncan have a different center point than the first positions and/or have afraction of a pixel pitch movement increment. And this step maybecombined with the earlier step by predetermining which one or moreregions of the XY area, movement increment size. These regions then havehigher density of spots of x ray emission positions. The order ofmovement, to the first positions or to the second positions, in someapplications, is not critical, and may be done in one scan as they maybe interwoven in the selected area.

The present disclosure also provides methods in which, independent of 3Dimaging measurements, if the relative spatial position of the region ofinterest, or the subject to the x ray source changes in any 6D dimension(for example by the movement or the rotation of the region of interestor by movement or the rotation of the x ray source) the system can usethe movement to resolve a 3D image. Because of the movement in 6D,elements of the detector will correspond to a different projected pathin the region of interest. The movement may be divided in increments ofat least one measurement unit with different x ray measurements for eachspatial position of the subject during the movement. The geometry of themovement relative to the x ray source and detector may be calibratedspatially. Each projected path detected by a corresponding detectorelement or pixel or element region of the detector with two or morepixels may serve as a data point in solving for the unknown pixels orvoxels in the Z axis for the region of interest. The newly introducedunknowns because of the movements will be determined on a case by casebasis.

A system where the x ray emitting position is moved in an xy planeparallel to the detector may be combined with the above x raymeasurements during such movement.

Such 3D imaging method is fast and simplified.

However when such a system is not available, in one configuration ofthis disclosure to include multiple dimension tomography known to thosefamiliar to the art as an alternative step to the 3D imaging methoddescribed here as part of the present disclosure to improve otheralternative 3D tomography systems based on 2D detectors.

2D Imaging

Multidimensional images are generated from 2D images taken from at leasttwo different x-ray emitting locations. The following provides oneexample procedure for determining the 2D images using an x-ray machine.

A determination of the geometry and dimensions of the subject or theregion of interest in the subject can be determined. If such informationis predetermined or preset, this step can be skipped.

The subject is then illuminated with x-rays from a first x-ray emittinglocation. The image is then read at the detector assembly.

Next, move the x-ray emitting location to a second location in the XYplane parallel to the plane of the detector assembly, where thedisplacement from the first location to the second location is aninteger multiple of the detector pixel pitch. When the displacement isone pixel pitch, a projected image of the region of interest of thesubject differs from the previous projected image by extending the outeredge of the image of the region of interest by exactly one line ofpixels on the detector assembly along the axis of the change ofdirection. In another words, moving the x-ray source or emittinglocation generates two different projected images for a region ofinterest of defined dimensions, and the location of the projected imageson the detector for the same region of interest is only extended by onepixel pitch in the direction of shift.

Repeat the x-ray emitting location step and measurement step.

Alternatively, there are various ways to generate projected images onthe detector with the end goal of minimizing new unknown pixels outsideof the region of interest in the projected paths in the 3D imagingprocess. For example, the x-ray source emitting position may be modifiedto move in 3D space or in at least one dimension. As long as a certainaspect or a portion of imaging process, if not the complete imagingprocess, use the present method, the imaging methods of the prior artmay be improved. Alternatively, in cases where the subject or the regionof interest moves, for example, in a conveyer belt, or a movingcomponent in the region of interest, such as tracking an unknowncomponent inside a cavity, the movements varies illumination paths ofthe component and at the same time generate the new unknown pixels inthe illuminating paths of the component. Such movements alone may besufficient to complete a 3D imaging dataset. In some cases, in order tocomplete 3D imaging, the relative position of the x-ray emittingposition or x-ray source may still need to be moved in order to completethe 3D illumination and imaging dataset. The number of unknown pixelsmay be increased or varied. However, as long as the illumination pathsare varied enough to resolve the complete 3D volume in the region ofinterest and newly introduced unknown pixels, complete 3D imaging can beaccomplished. Repeat the x-ray measurements with as different locationsof the x-ray source as needed in order to produce a 3D image of thedesired resolution in the Z axis.

Primary X-Rays and Scatter Separation

In this step, scatter is separated from the primary x-rays for each ofthe 2D images acquired above. A number of different scatter removal orscatter and primary separation methods are described as above.

The present disclosure includes preferred methods of scatter and primaryx ray separation where multiple x-ray energy sources are used, andscatter and primary x-rays are separated at each energy level, methodsand details of material decomposition can be found in PCT applicationPCT/US19/14391, which is incorporated herein in its entirety.

The configuration of adjustable beam selectors in FIGS. 4-15 , FIG. 8 ,and FIG. 9 are used in applications where the movement or x ray emissionsource or the emitting location of the x ray moves out of the alignmentof holes in the beam selector for the primary x ray to pass through, thebeam selector means are either insensitive to the x ray emissionposition, can remain aligned or can be adjusted to align again with thex ray source as described above in between 2D x ray image acquisition ona per need basis.

2D Functional Imaging

Although functional imaging is described above as an optionalindependent step, it actually is performed as modifications to the 2Dimaging. Functional imaging is providing information in addition tolocation or 2D visualization taken with a single-energy x-ray sourcewith or without scatter and primary x-ray separation. Examples offunctional imaging methods and systems are described below. Each exampleis independent of the others and may be combined to provide moreinformation as needed for applications.

Material Decomposition and Different Material Imaging

This method can be done for 2D images or after reconstructedmultidimensional images or reconstructed 3D images. For example,material decomposition can be done on measured projected 2D orsynthesized 2D or multidimensional images, or 3D images at dual- ormultiple-energy or sometimes single levels. Decomposition is the processby which dual- or multiple-energy x-rays are used to quantitativelyanalyze and separate components in the subject based on atomic z anddensity or other x ray sensitive characteristics of component in theregion of interest.

In one configuration, the x-ray source emits two x-ray pulses from eachx-ray source location: a high-energy pulse at an average energy level H,followed by a low-energy pulse at an average energy level L. In anotherconfiguration, the x-ray source emits three x-ray pulses from each x-raysource location: a high-energy pulse at an average energy level H,followed by a medium-energy pulse at an average energy level M, followedby a low-energy pulse at an average energy level L. In eachconfiguration, each pulse has a single, essentially unchanging energyspectrum. In another configuration, four or more energy pulses areemitted from the x-ray source.

In one configuration, rather than the 2D detectors described above thatcannot discriminate between different energy levels, the detectorassembly employs energy-sensitive, photon-counting detectors. Thesedetectors may be used with a conventional x-ray source, or with a timeof flight x-ray source, such as a picosecond x-ray source, to collectprimary x-rays for densitometry and quantitative analysis and separationof images for different components with varied atomic z. With aconventional x-ray source, the energy-sensitive photon-counting detectorreplaces the front detector, rear detector, or both in the dual-detectorplus beam selector assembly to ensure primary x-ray and scatterseparation, while at the same time, allowing dual-, triple-, ormultiple-energy and spectrum-energy imaging and spectroscopy ofdifferent materials or components in the subject.

Details of material decomposition using dual or multiple energy sourcescan be found in PCT application PCT/US19/14391, which is incorporatedherein in its entirety.

The present disclosure also provides methods in which single energy xray measurements are used for material decomposition. The method canhave the following optional elements:

-   -   conducts multiple dimension or 3D X ray imaging of the region of        interest;    -   such images are of primary x ray image with scatter reduced to        less than 1% of the primary signal, or in some cases less than        5% or in some cases, less than 10%;    -   determines one or more estimated material characteristics, such        as atomic number and density, spatial position, and other        characteristics which are defined by temporal marker or anatomic        marker, or one of a marker defined by facts deduced from digital        analytical algorithms performed on data derived from multiple        sources of the same or similar region of interest in the same or        similar subjects;    -   for systems with multiple measurement units in the region of        interest, using a tomographic reconstruction method described in        the present disclosure (these estimated material characteristics        can then be modified by reference to stored known material        characteristic data);    -   determining the composition of the volume in the region of        interest during reconstruction includes segmenting one or more        regions of interest into components, each with a common        composition;    -   the segmenting can be performed during iterative reconstruction        instead of being based on the voxel characteristics determined        upon the completion of iterative reconstruction;    -   one or more additional iterations of the tomographic        reconstruction algorithm, where each iteration updates the one        or more estimated material characteristics for components in the        region of interest.

Functional imaging is improved when primary x-rays and scatter areseparated.

Material Separation and Imaging

Interferometry

In one configuration, an interferometer is employed with or withoutscatter removal. With this, 2D images of absorption, dark field, and/orphase contrast images can all be obtained. Such images are used toconstruct a 3D interferogram.

The interferometer operates by emitting x-rays through a phase gratingthat introduces an interference fringe at specific distances downstream.When a subject is placed in the beam's path, the subject modifies theobserved interference pattern via absorption, refraction, andsmall-angle scattering. Once these signals are read by the detector, theproperties of the subject and its components can be determinedalgorithmically. In one example, Talbot-Lau interferometry is used inorder to have a large field of view. In Talbot Lau interferometry, abeam splitter grating (G1) is placed in the beam path between an x-raysource (S) and detector (D). Due to the fractional Talbot effect, anintensity distribution (I) revealing the periodic structure of the beamsplitter grating occurs at certain distances behind the grating. If anobject (O) is placed in front of the beam splitter grating, theintensity distribution changes due to the absorbing, scattering, andrefractive characteristics of the object. The fractional Talbot effectrequires spatially coherent radiation. To meet this requirement, amicrofocus x-ray tube with a sufficiently small focal spot can be used.Alternatively, a slit mask (G0) can be placed in front of the focal spotof a conventional x-ray tube. The mask absorbs certain parts of thex-ray beam and thereby creates spatially coherent slit sources. Each ofthese slit sources generates a self-image of the beam splitter grating.By exploiting the Lau effect, it is ensured that these self-imagessuperimpose to a sharp intensity distribution. In general, theseinterference fringes are too small to be resolved by a conventionalx-ray detector. To overcome this challenge, an absorbing analyzergrating (G2) with the same period as the interference fringes is placedat the plane of these fringes. This analyzer grating is used to samplethe periodic intensity distribution by shifting it stepwise in its planeperpendicular to its grating bars.

In order to generate coherent x-ray beams, for example, theinterferometer uses a pixilated x-ray source or coherent source grating.In another example, the interferometer has a diffraction grating, whichis MEM-based, crystal-based, or employs an acoustic modulated crystalgrating.

3D Imaging

High-resolution 3D imaging of a subject with defined dimensions, has thefollowing steps.

Calibration is as described above.

Primary x-ray and scatter separation and/or functional imaging. Examplesare as described in U.S. Provisional Patent Application Nos. 62/620,158,62/628,370, and 62/628,351, as explained above, dual- andmultiple-energy imaging.

The x-ray emitting location shifts relative to the subject to a secondposition in an x-ray plane parallel to the plane of the detector. Theshift distance between the x-ray emitting positions is set so that eachsubsequent image contains a projected image of the region of interestwhose location differs from the previous image by extending the outeredge of the detected image for the region of interest by one line ofpixels on the detector, and along the axis of shifting direction. Repeatthe x-ray measurements for each x-ray emitting location. The locationscan also be shifted by less than one line of pixels, or by more than oneline of pixels.

Determine the geometry or the dimensions of region of interest in thesubject. If such information is predetermined and stored in thecomputing device, skip this step.

Based on the thickness in the Z axis perpendicular to the x-ray plane,determine the total number of pixel-wide x-ray-emitting locations Pneeded in order to produce a complete 3D image, whereinP=thickness/pixel pitch (or resolution needed for the z axis)=n²+4n. Ifsuch information is predetermined, skip this step.

Repeat moving the x-ray emitting positions to the first positions andx-ray measurement at least P times in the XY plane with the emittingpositions in a travel area limited by coordinates on the x- or y-axis:smaller or equal to √(n²+4n). In cases when 4n is sufficiently smallcompared to P, p-4n first positions are sufficient for 3d imagingreconstruction. To resolve unknowns introduced in the new x ray emissionposition, second position are scanned. This is when x ray emissionpositions are closer in distance to adjacent second positions than thoseof adjacent first positions, but travel in the same area as the firstpositions.

It is noted that for x ray with pixilation x ray source, instead ofmoving the x ray source or physically moving x ray emission position,different pixels are used as the x ray source, each time an image istaken.

Combine x-ray measurements as described above, solve and determine theunknown pixels in the Z axis for the region of interest in a linearequation system m×n×p. Solve for and determine the new unknown pixelscreated from all x-ray emitting locations other than the first location.

The computing device uses a conventional computing tomography imagingalgorithm, including simply plugging in the resolved unknown pixelvalues to derive a 3D image of the region of interest base on theprevious steps. The computing device provides a multiple-axisrepresentation at various resolutions of 2D images combined with amultidimensional representation of both for the region of interest inthe subject.

The x-ray emitting location can be moved in planar space less than 5mm×5 mm, 4 mm×4 mm, 3 mm×3 mm, 2 mm×2 mm, or 1 mm×1 mm to derivecomplete 3D images of 100 μm resolution of a chest image of 20 cm×20cm×20 cm. This allows for a significant increase in the speed ofmeasurements and a significant drop in the amount of radiation requiredto obtain a 3D image. As a result, patient safety and comfort isdramatically increased.

The present disclosure includes embodiments where the x-ray measurementsfor a subject of 20 cm×20 cm×20 cm to derive a complete 3D image usingless than 1150 images or in some instances 1000 images at 100 μmresolution, in some instances, with scatter being <10% or 5% or 1% ofthe primary x-ray signals. With higher resolution, such as 100 nm or 10nm, proportionally, the number of x-rays measured will be increased, forexample 1 million images. However, for a 10 mm×10 mm×10 mm sample, at100 nm resolution, there will be 100,000 measurements, or 30 μm×30 μmmovement of the x-ray emitting location in the XY plane to derive acomplete 3D image.

It is also one aspect of the present disclosure to image large subjects,such as 25 cm×25 cm×25 cm, with small movements in the x and ydirections, such as 5 mm or sometimes even less, for example, 1 mm,relative movement of the x-ray emitting locations to the subject isaccomplished on a two dimensional plane with resolution or distancebetween adjacent x ray emitting locations to be in the 100 μm, or 0.1 umdepending on the resolution required and achievable resolution, definedby the pixel pitch of the detector hardware. The furthest x ray emittingposition from the original position maybe less than 5 degrees, or 4degrees or 2 degrees or less than 1.5 degrees in order to achievecomplete 3D imaging with minimum number of 2d images taken. For lowerresolution measurements such as 500 um for a subject with 25 cm inthickness or in depth, the furthest x ray emitting position maybe 7.5 cmfrom the original position, with significantly less number of imagestaken than for the higher resolution such as 100 um or 0.1 um in orderto complete 3D image reconstruction

The present disclosure allows measurements and thereby 3D imageconstruction of a subject in the subnanometer range, such as 0.01 nm or0.1 nm. For example, for a subject with dimensions in 10 μm×10 μm×10 μm,10 μm being the thickness, using an x-ray microscopy, the resolution canbe in the subnanometer range. An x ray microscope, as illustrated inFIG. 31 (35), which can have a condenser lens m-1, aperture, m-2,objective lens, m-4, and an x-ray detector assembly, m-5. M-5 can be,for example, a x ray detector or photodiode arrays or photon counter,spectrometer comprising an energy dispersive x ray optics such as agrating, and/or a spatially sensitive detector. In some cases, ascintillator to convert the x-ray signal into optical light can use thefollowing to measure the signal: a camera sensor, photon sensitivecamera or PMTs or PMT array or The subject can be placed immediatelyafter the aperture. X-rays emitted from an undulator as insertion devicein an electron (or positron) storage ring is monochromatized, m-7, andfocused onto the sample, m-3 by a condenser, m-1, which can be a zoneplate, whereby the objective lens magnifies the signal onto a detectorm-5. The sample is raster scanned and the transmitted intensitymonitored by a x-ray sensitive detector m-5. The distance between themost adjacent x ray emitting positions can be as small as 0.01 nm to 100nm depending on the application need. The condenser lens, m-1, can alsobe a specialized X-ray optics called multilayer Laue lenses (MLLs) whichare two perpendicularly oriented lenses. These lenses consist ofalternating layers of two different materials with nanometer thickness.

With MLLs, the resolution of a x ray microscope can reach 1 nm. Inreality, with the present dis, the limitation of x ray microscopy is thex ray condenser optics, theoretically, the present disclosure enablesresolution of along the z axis to be the diffraction limit of the x raywavelength. However, the present limitation is the x ray optics forfocusing x ray into a spot as small as 0.01 nm.

A collimator maybe used to select regions of radiation to reduceradiation on the subject. Or alternatively, an anode target maybemodulated and rotated so that regions on the anode target maybe areselected for x ray generation. In another configuration, electron beamemitting source maybe modified to selectively activate electron emittingunit so that x ray radiation is only generated in selected regions. Forexample, in a field emitter based x ray source, only selected regions offield emitter are activated. For example if 200 elements of fieldemitter material, such as nanotubes or nanowires are needed to generateenough electron current for the generation of x ray needed for theapplication, one or more field emitter element, such as nanotubes, ornanowires, may form a field emitting unit, which can also include acomputing apparatus, and maybe be activated independently from oneanother. For example, each field emitter unit may have a similar numberof field emitting elements, such as illustrated in FIG. 24 , C7.However, in other configurations, each field emitting regions may havedifferent numbers of field emitting elements than the rest.

Mechanisms for activation of each field emitting unit may be the same orsimilar to what is known to those familiar with the art. A computerprocessor will dictate which region to be activated at what time. It isalso part of the present disclosure to multiplex activation of two ormore independent field emitting regions at different times.

Multiple Dimension or 3D Images

It is also an aspect of the present disclosure to include two or morex-ray emitting locations or two or more x-ray sources for themultidimensional imaging of two or more subjects in a region of interestusing the 3D imaging method as described. Each x-ray source or emittinglocation illuminates a distinct path in the region of interest on itsrespective subject. Such imaging may be synchronized. Selected regionmultidimensional imaging may be done at the same time with differentx-ray sources or x-ray emitting locations, but not synchronized witheach other. This means each x-ray emitting location may be moved toilluminate a different projection path on its designated subject or adifferent x-ray emitting location in the predetermined emitting locationis turned on for subsequent illumination. The result is that separate 2Dimages are acquired of different subjects in the region of interesteither at the same time synchronous or asynchronous or at differenttimes as required by the application.

Alternative to the above process of multiple dimension x-ray imagingmeasurements and reconstruction, another embodiment of 3D imageacquisition and generation is based on a scanning-beam digital x-ray(SBDX). SBDX uses an electromagnetically-scanned electron beam incidentupon a large-area transmission style tungsten target. The electron beamis raster scanned over a 2D array of source focal spot positions every1/15 of a second. A multi-hole collimator defines a series of narrowoverlapping x-ray beams convergent upon a 2D detector. The geometricrelationship among the narrow beam projections is constrained by theprecise and rigid geometry of the SBDX collimator and the fixed detectorposition. A typical SBDX system geometry is as follows:Source-detector-distance (SDD)=1500 mm; Source-axis-distance (SAD)=450mm; Focal spot positions=71×71; Focal spot pitch=2.3×2.3 mm; Nativedetector array=320×160; Native detector element pitch=0.33 mm; andDetector bin mode=2×2.

SBDX has an inherent tomosynthesis capability due to the use of inversegeometry beam scanning. A live display analogous to conventionalfluoroscopy is generated using a GPU-based real-time imagereconstructor. Each displayed 2D image frame is generated through atwo-stage reconstruction procedure. First, shift-and-add digitaltomosynthesis is performed to generate a stack of, for example, 32single-plane images with, for example, a 5 mm plane spacing. The pixelcenters for the stack of tomosynthesis images are defined such that afixed pixel position (for example, row 100, column 100) in the stackcorresponds to an x-ray originating at the detector center. Next, agradient filtering procedure is applied to each of the single-planeimages to identify local regions of high sharpness and contrast. Thefinal 2D “composite” image is then formed by selecting, for each pixelposition, the pixel value from the single-plane image with highestcontrast and sharpness. Due to the geometry of the tomosynthesis pixelcenters and the compositing procedure, the final composite image can beviewed as an inverted “virtual” cone-beam projection of the in-focusobjects in the subject volume. A virtual SBDX projection originates atthe center of the detector and falls on the source plane. The pitch ofthe virtual detector elements at the source plane is, for example, 0.23mm based on the set geometry.

Alternatively, various 3D computation tomography configurations andmethods can be utilized. These can include, for example, using amotorized x-ray source to span significant angles, move linearly 1D, 2D,3D or within 6D space, a pixilated x-ray source, or using methods otherthan what is described above for 3D reconstruction. These can all becombined with the primary x-ray and scatter separation methods similarto, derivatives of, or the same as described above to achieve theresolution required and quantitative x-ray tomography by separatescatter and primary x-rays so that the scatter is <10% or <5% or <1% ofthe primary x-ray signals.

In a system involving multidimensional x-ray imaging methods, whenscatter is small compared to the primary x-rays, for example in alow-scatter sample or small animal, multidimensional images may beconstructed without involving the primary x-ray and scatter separationstep.

It is an aspect of the present disclosure to include embodiments of amethod comprising of steps:

Deriving a low resolution 3D or 2D image as described above, using morethan one pixel or region of pixels as a unit of unknown voxel. Based onan image or images derived above, selecting a subject region to beimaged in higher resolution, for example, as high a resolution as asingle pixel. Higher resolutions may also be applied to methods wherevisible light, optical detectors are used downstream of an x-rayscintillator, converting x-rays to visible light, and captured by thevisible light, optical detectors or visible light photodiodes orphotodiode array. In this case, the pixel size resolution may be theresolution of the camera or visible light detector or photodiode.

A high-resolution image can be derived as described as above and alow-resolution image of a subject may be derived using methods describedabove to track the subject. Low-resolution may be defined with ameasurement unit of two or several pixels, rather than of one pixel. Forexample, for a 20-cm-thick subject, at a 5-pixel resolution and eachpixel at 100 μm in pixel pitch, each measurement unit is 5×100 μm=500μm. There would be 20 cm/500 μm=400 unknown units or 400 positions thatneed to be resolved. So only 20 unit×20 unit area of x-ray emittinglocations, or <4 mm×4 mm area where x-ray emitting locations need to bescanned in the XY plane.

Medical applications include, but are not limited to, detectingcardiovascular abnormalities, detecting factures and othermusculoskeletal injuries, aiding the diagnose of neurological diseases,dental, screening for cancers, diagnosis of thoracic complications andconditions, surgical and procedure guidance and biopsy guidance, andtreatment management and monitoring. Industrial applications include,but are not limited to, x-ray inspection and identification, security,and environmental issues.

It is one aspect of the present disclosure to include the followingmethods based on the existing hardware configuration due to the factthat the present disclosure enables detection of temporal markers anddynamic events in millisecond or faster in 2D and multiple dimensionsand 3D:

-   -   Time: time stamp each image taken, either 2D or 3D, or more        multiple dimensional, such as attach a DICOM label or adding a        time label.    -   Fluidic dynamics, flow direction, dynamic movement in 6D: by        using contrast agents such as x ray measurement sensitive        markers, nanoparticle labeled markers or microbubbles        administered in one or more phases. The administration maybe for        example, oral intake or by injection or inhaling or means known        to those familiar with the intake method of in vivo contrast        agents. The methods are to track, for example, blood flow,        liquid flow in a microfluidics or lab on a chip, tracking of a        component such as implant or surgical tool or biopsy probe, or        disease markers in regions of the body. 3D images or multiple        dimension images or gated x ray measurements are done at        discrete times during one or more time intervals of interest,        markers maybe located by multiple dimensional imaging or 2D        imaging or using spectral or multiple energy x ray decomposition        at the region of interest. This is useful when the component to        be tracked comprised of one or more regions which maybe be        distinguished in x ray measurements from the background, for        example, when such regions on the component is at an atomic z        significantly different from those materials in its background.    -   Colocation with other measurement modalities, such as optical        spectroscopy, endoscope, optical tomography, ultrasound,        electrophysiology, MRI or SPECT, or PET or other x ray based        measurements, by identification and/or localization of one or        more contrast agents shared by x ray and the modalities, one or        more spatial marker or anatomical markers or localization of one        or more substances both temporally or spatially. Such devices in        some configurations have an intracavity probe, such as a        catheter and guide wire. X ray imaging of the present disclosure        provides the methods to identify and guide the location of such        probe and devices by either using material decomposition method,        or by using a probe which has one or more regions that are        designed to be visible at designated x ray energy levels. In the        event, that different chemical or molecular markers are used, in        case of X ray markers or contrast agents are different from        those used for other modalities, x ray contrast agents maybe        conjugated with contrast agents for other modalities, or the x        ray contrast agents may relate to the contrast agents of other        modalities spatially.    -   3D imaging and x ray analysis of components inside of a region        of interest using chemical based pertubation, such as a drug or        energy perturbation such as ultrasound or electromagnetic waves        such as laser or radiofrequency means such as in RF ablation of        tissues such as heart or kidney, or varied ultrasound        perturbation of capillaries or other tissues, the changes in        tissue composition in ablated or affected regions may give        different x ray measurements during the energy perturbation. An        optical probe or ultrasound probe device or catheter which        delivers energy such as RF may be used. Such as probe may have        regions of material compositions different from the background,        which result in varied x ray measurements. Additional, using 2D        functional imaging and or 3D functional imaging methods,        including material decomposition methods, such changes may be        identified and localized. Databases of markers associated with        changes during each phases of perturbation maybe established        ahead of time to allow referencing and look up to help diagnosis        and identification of stages of perturbation.    -   Identification of the image or image set: for example, each        image is labeled with at least the name and or a description of        the subject or the region of interest, or at least a unique        identification number or binary identification number, or all of        the aforementioned id information.    -   Recording number of images taken per subject based on DICOM        labels or unique identifier for each imaging process or each        imaging session or each study or treatment or diagnostic or        monitoring or therapeutic planning, or research project or        tracking period.    -   Recording and tally number of images taken and or processed for        each x ray system including the computer, the x ray hardware and        the software; a memory storage unit, electronically store one or        more documents, each has reports or up to date records of number        of images taken during a time frame such as a day or a month or        a year or since the system has been in use; the report or the        document can be accessed by either physically accessing the        computer and its associated x ray imaging system or the        electronic memory storage unit remotely via internet or intranet        or direct physical access for example a memory stick or security        key capable of storing and process digital information; a        computer is programmed to generate a report based on the        document, store electronically, and periodically automatically        sends the report to the predetermined recipient via email or        hardcopy or other electronic means for example, stored on a        server, password protected, accessible for the predetermined        recipient who can access by login to access the record by using        a password either at the x ray system location, and or at a        remote location.

The above describes a Minimized Spatial Variance for between eachadjacent image acquired and the total deviation from the originalimaging position, Minimized Imaging Acquisition Time, Minimized inRadiation and Complexicity, Radiolography based CT (3MR-CT), itsvariant, Spectral or multiple energy version is described as3MR-SPECTRAL CT and its variant, functional imaging version isdescribing as 3MR-F-CT.

Thus it has been shown and described a 3D imaging method. Since certainchanges may be made in the present disclosure without departing from thescope of the present disclosure, it is intended that all matterdescribed in the foregoing specification and shown in the accompanyingdrawings be interpreted as illustrative and not in a limiting sense.

1.-91. (canceled)
 92. An x-ray measurement device capable of determiningthree dimensional x-ray images of a region of interest (ROI) in asubject, the x-ray measurement device comprising: an x-ray source, thex-ray source configured to emit x-ray radiation at the ROI, wherein aplurality of projection images are acquired with x-ray radiation emittedfrom a plurality of different x-ray emitting positions relative to thesubject, the number of different x-ray emitting positions beingquantitively related to a dimension of the ROI along an axis; an x-raydetector comprising a plurality of detector elements arranged in a twodimensional plane that is perpendicular to the axis, opposite the x-raysource, the x-ray detector configured to detect x-ray radiation afterattenuation of the x-ray radiation by the ROI of the subject and providean indication of the detected x-rays; and a processor configured toreceive the indication of the detected x-rays and reconstruct athree-dimensional x-ray image of the ROI from the detected x-rayradiation.
 93. An x-ray measurement device capable of determining threedimensional x-ray images of a region of interest (ROI) in a subject, thex-ray measurement device comprising: an x-ray source, the x-ray sourceconfigured to emit x-ray radiation at the ROI, wherein a plurality ofprojection images are acquired with x-ray radiation emitted from aplurality of different x-ray emitting positions relative to the subject,the number of different x-ray emitting positions being determined basedon a ratio of a dimension of the ROI along an axis to a resolution ofthe ROI along the axis; an x-ray detector comprising a plurality ofdetector elements arranged in a two dimensional plane that isperpendicular to the axis, opposite the x-ray source, the x-ray detectorconfigured to detect x-ray radiation after attenuation of the x-rayradiation by the ROI of the subject and provide an indication of thedetected x-rays; and a processor configured to receive the indication ofthe detected x-rays and reconstruct a three-dimensional x-ray image ofthe ROI from the detected x-ray radiation.